Microfabricated device with micro-environment sensors for assaying coagulation in fluid samples

ABSTRACT

The present invention relates to sample analysis cartridges comprising micro-environment sensors and methods for assaying coagulation in a fluid sample applied to the micro-environment sensors, and in particular, to performing coagulation assays using micro-environment sensors in a point of care sample analysis cartridge. For example, the present invention may be directed to a sample analysis cartridge including an inlet chamber configured to receive a biological sample, and a conduit fluidically connected to the inlet chamber and configured to receive the biological sample from the inlet chamber. The conduit may include a micro-environment prothrombin time (PT) sensor, and a micro-environment activated partial thromboplastin time (aPTT) sensor.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a divisional of U.S. Ser. No. 16/259,500, filed Jan.28, 2019, which is a divisional of U.S. Ser. No. 14/866,316, filed Sep.25, 2015, now U.S. Pat. No. 10,247,741, issued Apr. 2, 2019, whichclaims priority to U.S. Provisional Application No. 62/055,776 filedSep. 26, 2014, and U.S. Provisional Application No. 62/055,777 filedSep. 26, 2014. The entireties of the foregoing are incorporated hereinby reference.

FIELD OF THE INVENTION

The present invention relates to sample analysis cartridges comprisingmicro-environment sensors and methods for assaying coagulation in afluid sample applied to the micro-environment sensors, and inparticular, to performing coagulation assays using micro-environmentsensors in a point of care sample analysis cartridge.

BACKGROUND OF THE INVENTION

Blood clotting or hemostasis is an important protective mechanism of thebody for sealing wounds caused from injury to the body. Hemostasis takesplace in two phases. Primary (cellular) hemostasis serves to quicklystop bleeding and minimize blood loss. Primary hemostasis involvesinjured cells of the endothelium and the underlying layer of cellsemitting signals that enable blood platelets (thrombocytes) toaccumulate in a region of an injured blood vessel, forming a plug thatprovisionally seals the wound. Secondary (plasmatic) hemostasis orcoagulation is initiated at the same time as primary hemostasis andinvolves a process by which blood clots. More specifically, coagulationis controlled by a signaling coagulation cascade consisting of thirteencoagulation factors that interact and activate each other. At the end ofthe coagulation cascade, fibrinogen is converted into fibrin. A networkof fibrin fibers reinforces wound closure, and platelets and other bloodcells get caught in this network and form a blood clot (thrombus).Lastly, platelets and the endothelium release growth factors thatcontrol a wound-healing process. At the end of these processes, thefibrin network is dissolved by enzymes in the blood plasma.

Hemostasis requires a subtle balance of procoagulants and anticoagulantssuch that circulating blood remains a relatively low-viscosity fluid andcoagulation only begins in order to seal wounds. Procoagulants preventexcessive bleeding by blocking blood flow from a wound or damagedvessel, whereas anticoagulants prevent clots from forming in thecirculating system, which could otherwise block blood vessels and leadto myocardial infarction or stroke.

The coagulation cascade of secondary hemostasis is based on catalyticconversion of fibrinogen, a soluble plasma protein, to insoluble fibrin.The enzyme catalyzing this reaction is thrombin, which does notpermanently circulate in the blood in an active form but exists asprothrombin, the inactive precursor of thrombin. The coagulation cascadeleading to active thrombin consists of two pathways, the extrinsic andthe intrinsic pathways, which converge into a common pathway thatincludes active thrombin catalyzing the conversion of fibrinogen tofibrin. The extrinsic pathway is initiated at the site of injury inresponse to the release of tissue factor (factor III) and thus, is alsoknown as the tissue factor pathway. Tissue factor is a cofactor in thefactor VIIa-catalyzed activation of factor X (inactive) to factor Xa(active). The second, more complex, intrinsic pathway is activated byclotting factors VIII, IX, X, XI, and XII associated with platelets.Also required are the proteins prekallikrein (PK) andhigh-molecular-weight kininogen (HK or HMWK), as well as calcium ionsand phospholipids secreted from platelets. Each of these constituentsleads to the conversion of factor X to factor Xa. The common point inboth pathways is the activation of factor X to factor Xa. Factor Xa isan enzyme (e.g., a serine endopeptidase) that cleaves prothrombin in twoplaces (an arg-thr and then an arg-ile bond), which yields activethrombin and ultimately results in the conversion of fibrinogen tofibrin.

Breakdown of a blood clot or the fibrin network, termed fibrinolysis,requires the conversion of fibrin to a soluble product. This lysis iscatalyzed by the proteolytic enzyme plasmin, which circulates in aninactive form, plasminogen. Tissue plasminogen activator (tPA),bacterial hemolytic enzymes (e.g., streptokinase), and proteolytic humanenzymes found in urine (e.g., urokinase) all activate plasminogen. Thesematerials are typically used in thrombolytic therapy.

Consequently, the coagulation cascade is a suitable target fordiagnosing and treating diseases involving dysregulated blood clottingor the absence of clotting. For example, the diagnosis of hemorrhagicconditions such as hemophilia, where one or more of the thirteen bloodclotting factors involved in the coagulation cascade may be defective,can be achieved by a wide variety of coagulation tests. In addition,several tests have been developed to monitor the progress ofthrombolytic therapy. Other tests have been developed to signal aprethrombolytic or hypercoagulable state, or monitor the effect ofadministering protamine to patients during cardiopulmonary bypasssurgery. However, the main value of coagulation tests is in monitoringoral and intravenous anticoagulation therapy. Three of the keydiagnostic tests are prothrombin time (PT), activated partialthromboplastin time (aPTT), and activated clotting time (ACT).

PT is the time it takes plasma to clot after the addition of tissuefactor (obtained from animals such as rabbits, or recombinant tissuefactor, or from brains of autopsy patients). This measures the qualityof the extrinsic pathway (as well as the common pathway) of coagulation.The PT is most commonly used to monitor oral anticoagulation therapy.Oral anticoagulants such as Coumadin® suppress the formation ofprothrombin. The traditional PT test includes blood being drawn into atube containing liquid sodium citrate, which acts as an anticoagulant bybinding the calcium in a sample. Consequently, the PT test is based onthe addition of calcium and tissue thromboplastin to the citrated bloodsample, and the time the sample takes to clot is measured.

aPTT is the time taken for a fibrin clot to form. This measures thequality of the intrinsic pathway (as well as the common pathway) ofcoagulation. The aPTT is most commonly used to monitor intravenousheparin anticoagulation therapy. Heparin administration has the effectof suppressing clot formation. The traditional aPTT test includes bloodbeing drawn into a tube containing liquid sodium citrate, which acts asan anticoagulant by binding the calcium in a sample. Consequently, theaPTT test is based on the addition of activating agent, calcium, and aphospholipid to the citrated blood sample (e.g., a platelet poorplasma), and the time the sample takes to form a fibrin clot ismeasured.

ACT is the time taken for whole blood to clot upon exposure to anactivator. The intrinsic pathway test evaluates the intrinsic and commonpathways of coagulation. The ACT is most commonly used to monitor theeffect of high-dose heparin before, during, and shortly after proceduresthat require intense anticoagulant administration, such as cardiacbypass, cardiac angioplasty, thrombolysis, extra-corporeal membraneoxygenation (ECMO) and continuous dialysis. The traditional ACT testincludes whole blood being added to a tube containing a surfaceactivator (e.g., celite, kaolin, or glass balls), which results in theactivation of the coagulation cascade via the intrinsic (Factor XII)pathway. Consequently, the ACT test is based on the addition of anactivator to the intrinsic pathway to fresh whole blood to which noexogenous anticoagulant has been added, and the time the sample takes toform a fibrin clot is measured.

Coagulation monitors are known for the analysis of whole blood. Forexample, a capillary flow device has been described in U.S. Pat. No.4,756,884 in which dry reagents are placed into an analyzer, which isthen heated to 37° C. before a drop of blood is introduced. The sampleis mixed with the reagent by capillary draw. The detection mechanism isbased on laser light passing through the sample. Blood cells movingalong the flow path yield a speckled pattern specific to unclottedblood. When the blood clots, movement ceases producing a patternspecific to clotted blood. A bibulous matrix with dried coagulationreagents has been devised for a single coagulation test in a device(See, e.g., U.S. Pat. No. 5,344,754) with integrated means fordetermining a change in electrical resistance upon addition of a sampleto the matrix. Detection of the reaction is based on a separate opticalassembly that is aligned with and interrogates the bibulous region ofthe device.

Coagulation point of care assays are also known for the analysis offluid samples or biological samples. For example, point of carecartridges for conducting a variety of assays responsive to a change inthe viscosity of a fluid sample, including assays involving whole bloodcoagulation, agglutination, fibrinolysis tests and, generally, assaysfor obtaining information on the clotting or lytic (lysis) process areknown (See, for example, U.S. Pat. Nos. 5,447,440 and 5,628,961, whichare incorporated herein by reference in their entireties). Additionally,point of care cartridges that provide a means by which a blood samplecan be metered and quantitatively mixed with reagents that activate theprimary or secondary pathway of the coagulation cascade for subsequentdetection of clot formation using a microfabricated sensor are known(See, for example, U.S. Pat. Nos. 6,750,053; 7,923,256; 7,977,106 and6,438,498, which are incorporated herein by reference in theirentireties).

However, coagulation point of care assay systems configured to performthe aforementioned coagulation assays of fluid samples generallycomprise the reagent and substrate printed in a dissolvable form on acover or base of the point of care cartridge or testing device. Duringanalysis, the sample is pushed and pulled by a mechanical process todissolve and mix the reagent and substrate into the sample. Thisarrangement of having the reagent and substrate printed in this form incombination with the requirement for mixing the reagent and substrateinto the sample has hindered the integration of coagulation tests into asingle point of care cartridge or testing device because of thepotential for cross-activation of the two distinct coagulation cascadepathways. Accordingly, the need exists for improved point of carecartridge or testing device design that allows for a combination ofcoagulation tests to be performed on a single point of care cartridge ortesting device.

SUMMARY OF THE INVENTION

In one embodiment, the present invention is directed to a sampleanalysis cartridge including an inlet chamber configured to receive abiological sample, and a conduit fluidically connected to the inletchamber and configured to receive the biological sample from the inletchamber. The conduit includes a micro-environment PT sensor and amicro-environment aPTT sensor.

Optionally, the sample analysis cartridge may further include a pumpconfigured to move a biological sample in the conduit and position thebiological sample over the micro-environment PT sensor, and move thebiological sample in the conduit and position the biological sample overthe micro-environment aPTT sensor.

In another embodiment, the present invention is directed to a sampleanalysis cartridge including an inlet chamber configured to receive abiological sample, a first conduit fluidically connected to the inletchamber and configured to receive the biological sample from the inletchamber, the first conduit including a junction configured to split thebiological sample into at least first and second segments, a secondconduit fluidically connected to the first conduit at the junction andconfigured to receive the first segment of the biological sample, thesecond conduit including PT reagent and a PT sensor, and a third conduitfluidically connected to the first conduit at the junction andconfigured to receive the second segment of the biological sample, thethird conduit including an aPTT reagent and an aPTT sensor. One of thesecond and third conduits includes a flow restrictor positioned upstreamof the PT sensor or the aPTT sensor.

Optionally, the sample analysis cartridge may further include a pumpconfigured to move the first segment in the second conduit to dissolvethe PT reagent into the first segment and position the first segmentover the PT sensor, and move the second segment in the third conduit todissolve the aPTT reagent into the second segment and position thesecond segment over the aPTT sensor.

In another embodiment, the present invention is directed to a sampleanalysis cartridge including a conduit including a micro-environmentcoagulation sensor and a blood chemistry sensor, and a pump configuredto move a biological sample within the conduit to contact themicro-environment coagulation sensor and the blood chemistry sensor atsubstantially a same time.

In yet another embodiment, the present invention is directed to a sampleanalysis cartridge including a conduit including at least onecoagulation sensor, a pH sensor, a partial pressure CO₂ sensor, apartial pressure O₂ sensor, and a lactate sensor, and a pump configuredto move a biological sample within the conduit to contact each of the atleast one coagulation sensor, the pH sensor, the partial pressure CO₂sensor, the partial pressure O₂ sensor, and the lactate sensor.

In some embodiments, the present invention is directed to a chipincluding a micro-environment PT sensor, a micro-environment aPTTsensor, and a counter reference electrode. Optionally, the PT and theaPTT sensors are electrochemical sensors.

In some embodiments, the chip has a substantially planar surface.Optionally, the counter reference electrode is a silver/silver chlorideelectrode.

In alternative embodiments, the present invention is directed to a chipincluding a micro-environment PT sensor, a micro-environment ACT sensor,and a counter reference electrode. Optionally, the PT and the ACTsensors are electrochemical sensors.

In some embodiments, the chip has a substantially planar surface.Optionally, the counter reference electrode is a silver/silver chlorideelectrode.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention will be better understood in view of the followingnon-limiting figures, in which:

FIG. 1 shows a cartridge schematic in accordance with some aspects ofthe invention;

FIGS. 2 and 3 show a conduit comprising a dissolvable reagent/substrateand transducer in accordance with some aspects of the invention;

FIG. 4 shows a diffusible reagent, immobilized substrate-polymer layer,and transducer in accordance with some aspects of the present invention;

FIGS. 5 and 6 show graphs that provide empirical evidence for aspects ofthe present invention;

FIGS. 7A, 7B, and 7C illustrate the principle of operation of themicroenvironment sensor comprising a reagent and/or substrate,immobilized or not in a polymer layer, and transducer in accordance withsome aspects of the invention;

FIG. 8 shows a cartridge schematic in accordance with some aspects ofthe invention;

FIG. 9 shows a side view of the fabrication of an immobilizedreagent/substrate-polymer layer in accordance with some aspects of theinvention;

FIGS. 10-12 show graphs that provide empirical evidence for aspects ofthe present invention;

FIGS. 13A, 13B, and 13C show multiple arrangements for a diffusiblereagent, immobilized substrate-polymer layer, and transducer inaccordance with some aspects of the invention;

FIG. 14 shows a side view of the fabrication of a sensor in accordancewith some aspects of the invention;

FIGS. 15 and 16 show multiple sensor configurations in accordance withsome aspects of the invention;

FIG. 17 shows a top view of a disposable sensing device in accordancewith some aspects of the invention;

FIGS. 18-20 show multiple sensor configurations in accordance with someaspects of the invention;

FIGS. 21, 22A, and 22B illustrate the principle of operation forconductometric sensors in accordance with some aspects of the invention;

FIG. 23 shows an isometric view of a disposable sensing device andreader device in accordance with some aspects of the invention;

FIG. 24 shows a top view of a disposable sensing device in accordancewith some aspects of the invention;

FIGS. 25 and 26 show a top view of a portion of disposable sensingdevices in accordance with some aspects of the invention;

FIGS. 27-29 show advanced microfluidic systems in accordance with someaspects of the invention;

FIG. 30 shows a graph of independent mixing control in accordance withaspects of the invention;

FIG. 31 shows a top view of a portion of a disposable sensing device inaccordance with some aspects of the invention;

FIGS. 32 and 33 show advanced microfluidic systems in accordance withsome aspects of the invention;

FIG. 34 shows a top view of a portion of a disposable sensing device inaccordance with some aspects of the invention;

FIG. 35 shows an advanced microfluidic system in accordance with someaspects of the invention;

FIG. 36 shows a top view of a portion of a disposable sensing device inaccordance with some aspects of the invention;

FIGS. 37 and 38 show advanced microfluidic systems in accordance withsome aspects of the invention;

FIG. 39 shows a top view of a portion of a disposable sensing device inaccordance with some aspects of the invention;

FIGS. 40, 41A, and 41B show graphs that provide empirical evidence foraspects of the present invention; and

FIGS. 42 and 43 illustrate the principle of operation for eliminatingthe ground chip in accordance with some aspects of the invention.

DETAILED DESCRIPTION OF THE INVENTION

The present invention relates to sample analysis cartridges comprisingmicro-environment sensors and methods for assaying coagulation in afluid sample applied to the micro-environment sensors, and inparticular, to performing coagulation assays using micro-environmentsensors in a point of care sample analysis cartridge.

In some embodiments, the invention relates to an integrated circuit chiphaving one or more test sensors comprising at least one transducercoated with a polymer layer that includes a thrombin-cleavable peptidewith a detectable moiety such that the one or more sensors operate in alocalized manner and are capable of determining one or more diagnosticclotting times (e.g., PT, aPTT, and/or ACT). More specifically, theinvention relates to a sample analysis cartridge comprising an inletchamber configured to receive a biological sample (e.g., blood, plasma,serum, urine and modified and diluted forms thereof) and a conduitfluidically connected to the inlet chamber and configured to receive thebiological sample from the inlet chamber. The conduit may comprise afirst micro-environment sensor and a second micro-environment sensorthat are configured to operate in a localized manner and are capable ofdetermining, respectively, a first diagnostic clotting time (e.g., PT)and a second diagnostic clotting time (e.g., aPTT) different from thefirst diagnostic clotting time.

In some embodiments, the first micro-environment sensor may include atleast one transducer coated with a substantially heparin-neutralizingpolymer layer and a thrombin-cleavable peptide with a signal moiety. Insome embodiments, the second micro-environment sensor may include atleast one transducer coated with a substantiallynon-heparin-neutralizing polymer layer and a thrombin-cleavable peptidewith a signal moiety. The first and second micro-environment sensors mayfurther include, respectively, first and second diagnostic clotting timereagents within the polymer layers (e.g., the reagents are integratedwithin the polymer layers), coated over the polymer layers (e.g., thereagents are a separate layer dispensed on top of the polymer layers),or positioned substantially adjacent to the polymer layers and/or the atleast one transducer (e.g., the reagents are positioned within theconduit such that the reagents are abutted to or within an interactivedistance of the polymer layers and/or the at least one transducer so asto still function in conjunction with each other).

Additionally, the invention relates to advanced microfluidic systems forcontrol of the biological sample within the sample analysis cartridge.In preferable embodiments, the sample analysis cartridge design enablestwo physically separated tests (e.g., PT and aPTT) to be conductedsimultaneously on a single biological (e.g., whole blood) sample withinthe same sample analysis cartridge. In some embodiments, the advancedmicrofluidic systems may comprise passive fluidic features (e.g.,valves, resistances, and fluidic locking elements) in addition to activefluidic features from the analyzer (e.g., a pump) to split thebiological sample into separate conduits/regions of a sample analysiscartridge such that each sample segment can subsequently be moved to aspecific sensor (e.g., biosensor or micro-environment sensor asdiscussed in detail herein). In additional or alternative embodiments,an integrated circuit chip may comprise multi-conduit conductometricelectrodes (e.g., hematocrit bars) configured to provide multiple pointsof contact with the biological sample for advanced microfluidic controlover the sensors or micro-environment sensors.

As used herein, the term “micro-environment sensor” refers to a sensorconfigured such that any reaction occurring in the immediate vicinity ofthe sensor in a manner sufficient to achieve the desired signal at thesensor will not detectably interfere with (or impact) another reactionoccurring at an adjacent sensor during normal usage.

As used herein, the term “heparin neutralizing” refers to an aspect ofthe sensor which renders unfractionated heparin and low-molecular-weightheparin (LMWH) biologically inactive in a biological sample in an areasufficient to span the micro-environment sensor area. Conversely,“non-heparin-neutralizing” refers to an aspect of the sensor that doesnot impact/affect the biological activity of unfractionated heparin orLMWH in the micro-environment sensor area.

As used herein, the term “immobilized” refers to an aspect of themicro-environment sensor which is substantially limited in movement, andthus localizing this aspect of the micro-environment to a general area.

As used herein, the term “substrate” refers to either a molecule whichis the target of an enzymatic reaction or a physical entity which formsthe foundation of a structure.

Overview of Blood Coagulation

The process of blood clotting and the subsequent dissolution of the clotfollowing repair of the injured tissue is termed hemostasis. In orderfor hemostasis to occur, platelets must adhere to exposed collagen,release the contents of their granules, and aggregate. The adhesion ofplatelets to the collagen exposed on endothelial cell surfaces ismediated by von Willebrand factor (vWF). The activation of platelets viathrombin is required for their consequent aggregation to a plateletplug. However, equally significant is the role of activated plateletsurface phospholipids in the activation of the coagulation cascade.

The intrinsic pathway of the coagulation cascade requires the clottingfactors VIII, IX, X, XI, and XII. Also required are the proteinsprekallikrein (PK) and high-molecular-weight kininogen (HK or HMWK), aswell as calcium ions and phospholipids secreted from platelets. Each ofthese intrinsic pathway constituents leads to the conversion of factor Xto factor Xa. Initiation of the intrinsic pathway occurs whenprekallikrein, high-molecular-weight kininogen, factor XI and factor XIIare exposed to a negatively charged surface. This is termed the contactphase and can occur as a result of interaction with the phospholipids(primarily phosphatidylethanolamine, PE) of circulating lipoproteinparticles such as chylomicrons, very low density lipoproteins (VLDLs),and oxidized low density lipoproteins (LDLs). This is the basis of therole of hyperlipidemia in the promotion of a pro-thrombotic state.

The activation of factor Xa in the intrinsic pathway requires assemblageof the tenase complex (Ca²⁺ and factors VIIIa, IXa and X) on the surfaceof activated platelets. One of the responses of platelets to activationis the presentation of phosphatidylserine (PS) and phosphatidylinositol(PI) on their surfaces. The exposure of these phospholipids allows thetenase complex to form and the subsequent activation of factor Xa.

The extrinsic pathway of the coagulation cascade is initiated at thesite of injury in response to the release of tissue factor (factor III)and thus, is also known as the tissue factor pathway. Tissue factor is acofactor in the factor VIIa-catalyzed activation of factor X. FactorVIIa, a gla residue containing serine protease, cleaves factor X tofactor Xa in a manner identical to that of factor IXa of the intrinsicpathway. The activation of factor VII occurs through the action ofthrombin or factor Xa. The ability of factor Xa to activate factor VIIcreates a link between the intrinsic and extrinsic pathways.

The common point in both pathways is the activation of factor X tofactor Xa. Factor Xa activates prothrombin (factor II) to thrombin(factor IIa). Thrombin, in turn, converts fibrinogen to fibrin. Theactivation of thrombin occurs on the surface of activated platelets andrequires formation of a prothrombinase complex. This complex is composedof the platelet phospholipids, phosphatidylinositol andphosphatidylserine, Ca²⁺, factors Va and Xa, and prothrombin. Factor Vis a cofactor in the formation of the prothrombinase complex, similar tothe role of factor VIII in the tenase complex formation. Like factorVIII activation, factor V is activated to factor Va by means of minuteamounts and is inactivated by increased levels of thrombin. Factor Vabinds to specific receptors on the surfaces of activated platelets andforms a complex with prothrombin and factor Xa.

Prothrombin is a 72 kDa, single-chain protein containing ten glaresidues in its N-terminal region. Within the prothrombinase complex,prothrombin is cleaved at 2 sites by factor Xa. This cleavage generatesa 2-chain active thrombin molecule containing an A and a B chain whichare held together by a single disulfide bond. Thrombin binds to a classof G-protein-coupled receptors (GPCRs) called protease activatedreceptors (PARs), specifically PAR-1, -3 and -4. PARs utilize a uniquemechanism to convert the result of extracellular proteolytic cleavageinto an intracellular signaling event. PARs carry their own ligand,which remains inactive until protease cleavage, such as by thrombin,“unmasks” the ligand. Following thrombin cleavage the unmasked ligand isstill a part of the intact PAR but is now capable of interacting withthe ligand-binding domain of the PAR resulting in the activation ofnumerous signaling cascades.

Overview of Coagulation Testing

Bleeding time assays are used to evaluate the vascular and plateletresponses that are associated with hemostasis. The bleeding time is afrequent assay performed on preoperative patients to ensure there is anadequate response to vessel injury prior to surgery. As discussedherein, the rapid responses to vascular injury (occurring withinseconds) are vessel constriction and platelet adhesion to the vesselwall. The Ivy method for determining the bleeding time involves the useof a blood pressure cuff (sphygmomanometer) which is placed on theforearm and inflated to 40 mm Hg. A superficial incision is then made onthe forearm and the time it takes for bleeding to stop is recorded. Withthe Ivy method bleeding should stop within 1-9 minutes. Any bleedingtime greater than 15 minutes would be indicative of a defect in theinitial responses of vessels and platelets to vascular injury. A lessinvasive bleeding time assay involves the use of a lancet or specialneedle, with which a 3-4 mm deep prick is made on the fingertip orearlobe. This bleeding time assay is referred to as the Duke method, andin this assay bleeding should cease within 1-3 minutes. The bleedingtime is affected (prolonged) by any defect in platelet function, byvascular disorders, and in von Willebrand disease but is not affected byother coagulation factors. Disorders that are commonly associated withan increased bleeding time include thrombocytopenia, disseminatedintravascular coagulation (DIC), Bernard-Soulier syndrome and Glanzmannthrombasthenia. Abnormal bleeding times are also found in patients withCushing syndrome, severe liver disease, leukemia, and bone marrowfailure.

Defects associated with factors of the pathways of blood coagulation canalso be assessed with specific assays. The prothrombin time (PT) is anassay designed to screen for defects in fibrinogen, prothrombin, andfactors II, V, VII, and X and thus measures activities of the extrinsicpathway of coagulation. When any of these factors is deficient then thePT is prolonged. A normal PT is 11.0-12.5 seconds. A PT greater than 20seconds is indicative of coagulation deficit. The PT is commonlymeasured using plasma after the blood cells are removed. A blood sampleis typically collected in a tube containing citrate to bind any calciumand thus inhibit coagulation, and then the cells are separated bycentrifugation. Excess calcium is added to an aliquot of the plasma toinitiate coagulation. The most common measure of PT is to divide thetime of coagulation of a patient's blood by that of the mean normal PTvalue, with this ratio subsequently being raised to a powercorresponding to the ISI (international sensitivity index) of thereagent being used. The resulting value is referred to as theinternational normalized ratio (INR). Normal values range from 0.8-1.2INR. PT is used to determine the correct dosage of the coumarin class ofanti-coagulation drugs (e.g. Coumadin®), for the presence of liverdisease or damage, and to evaluate vitamin K status.

The activated partial thromboplastin time (aPTT) is used to assay fordefects in the intrinsic pathway of coagulation. The aPTT assay includesthe addition of activators that shorten the normal clotting time and isnormally prescribed in patients with unexplained bleeding or clotting.The assay will evaluate the function of fibrinogen, prothrombin, andfactors V, VIII, IX, X, XI, and XII. A defect in any of these factorswill result in a prolonged aPTT. A normal aPTT is 30-40 seconds. TheaPTT is a standard assay used to assess the efficacy of heparinanticoagulant therapy. The aPTT is commonly measured using plasma afterthe blood cells are removed. A blood sample is typically collected in atube containing citrate to bind any calcium and thus inhibitcoagulation, and then the cells are separated by centrifugation. Excesscalcium is added to an aliquot of the plasma to reverse citrateanticoagulation. Prolonged aPTTs are associated with acquired orcongenital bleeding disorders associated with coagulation factordeficiency, vitamin K deficiency, liver disease, DIC, von Willebranddisease, leukemia, hemophilia, and during heparin administration.

The activated clotting time (ACT) is a common point-of-care whole-bloodclotting test used to monitor high-dose heparin therapy or treatmentwith bivalirudin. The dose of heparin or bivalirudin required in thesesettings is beyond the range that can be measured with the aPTT.Typically, whole blood is collected into a tube or cartridge containinga coagulation activator (e.g., celite, kaolin, or glass particles) and amagnetic stir bar, and the time taken for the blood to clot is thenmeasured. The reference value for the ACT typically ranges between 70and 180 seconds. The desirable range for anticoagulation depends on theindication and the test method used. For example, during cardiopulmonarybypass surgery, the desired ACT range with heparin may exceed 400 to 500seconds. In contrast, in patients undergoing percutaneous coronaryinterventions, a target ACT of 200 seconds is advocated when heparin isadministered in conjunction with a glycoprotein IIb/IIIa antagonist,whereas an ACT between 250 and 350 seconds is targeted in the absence ofsuch adjunctive therapy.

Electrochemical System for the Determination of Diagnostic ClottingTimes

Chromogenic assays have been used to measure the enzymatic activity ofspecific clotting factors through the development of artificial,cleavable peptide substrates specific for particular factors. It shouldbe noted that assays based on clotting time, such as aPTT, PT and ACT,are essentially functional measures of thrombin formation and inhibitionin the presence of anticoagulants, such as warfarin and heparin ordefective coagulation factors. Thus, an analogy can be drawn betweenassays based on the measurement of fibrin formation and assays baseddirectly on the measurement of thrombin activity via the use ofappropriate peptide substrates, as in chromogenic assays.

Electrochemical detection involves the use of a working electrode (e.g.,an amperometric electrode) and a reference electrode (e.g., a counterreference electrode), whereby a constant potential is applied to theworking electrode leading to an oxidation-reduction (redox) reactionthat can be quantified as a recordable electric current. Electrochemicalsensors have found widespread use in the development of point-of-care(POC) and self-test devices, as exemplified by the development ofglucose test strips, as they are simple to interface with electronicinstruments and reduce device costs. Devices, such as the i-STAT® system(see, e.g., U.S. Pat. No. 7,977,106, the entirety of which isincorporated herein by reference), have employed electrogenic substratesthat result in the formation of an electrochemically detectable cleavageproduct that is proportional to thrombin activity. These devices arethen configured to return a clotting time based on a measure of thrombinactivity to allow comparisons with standard clotting. Accordingly, insome embodiments, the electrochemical detection system is termed“electrogenic” because the electrochemically detectable species aregenerated to allow determination of a rate measurement or a testendpoint, e.g., a diagnostic clotting time. This is similar to thechromogenic or fluorogenic endpoint tests in which a change in the lightabsorbing or emitting properties of a sample indicates the ratemeasurement or endpoint, e.g., a diagnostic clotting time.

FIG. 1 illustrates the principle of an electrochemical detection system10 (e.g., an amperometric electrochemical detection system) according tosome embodiments of the present invention for determination ofdiagnostic clotting times. However, it should be understood that whilespecific embodiments are described herein for diagnostic clotting timeassays (e.g., PT, aPTT, and ACT assays), the micro-environment sensorstructures described herein may also be useful for detecting variousanalytes of potential interest. More specifically, the electrochemicaldetection system of the present invention is not limited to the assay ofcoagulation enzymes. For example, any assay where an enzyme cleaves asubstrate molecule to yield an electroactive moiety can use the presentmethodology. As should be understood, assays can be devised for avariety of other known enzymes in the art, such as for example, glucoseoxidase, lactate oxidase, and other oxidoreductases, dehydrogenase basedenzymes, and alkaline phosphatase and other phosphatases, and serineproteases without departing from scope of the present invention. Forexample, some aspects of the present invention may include a phosphataseassay where ferrocene with a phosphate moiety is present in amicro-environment sensor layer. The enzyme phosphatase present in asample may permeate the micro-environment sensor and cleave thephosphate groups enabling the liberated ferrocene molecules to beoxidized at the electrode. Accordingly, the measured current may be afunction of the rate of the cleavage reaction, and thus, proportional tothe phosphatase activity in the sample.

In an exemplary analysis, a fluidic sample 15, e.g., whole blood, may beintroduced into a sample holding chamber 20 of a cartridge 25 of thepresent invention. Thereafter, the fluidic sample 15 may be introducedto an analysis region 30 of the cartridge, e.g., a sensor region or oneor more locations within one or more conduits of the cartridge thatincludes one or more sensors for coagulation detection and optionallyfor detection of a target analyte (e.g., thrombin activity for aprothrombin time and troponin I). The analysis region 30 includes one ormore micro-environment sensors 35 comprising one or more electrodes ortransducers 37, one or more reagents 40, and one or more substrates 45in any number of different possible arrangements. The form andorientation of the electrodes, reagents, and substrate may vary widelydepending on the embodiment of the invention, which are described indetail hereafter.

In accordance with some aspects of the invention, the one or morereagents 40 may include a material for inducing coagulation via theintrinsic or extrinsic pathway. Materials suitable for inducing theextrinsic pathway (e.g., PT analysis) may include one or more componentsselected from the group consisting of non-recombinant tissue factor,recombinant tissue factor, a synthetic or natural lipid, a synthetic ornatural phospholipid, a combination of synthetic or natural lipids, anda combination of synthetic or natural phospholipids. In some embodimentsa variety of other components may be included within the one or morereagents 40 to contribute to stabilization and deposition/dissolutioncharacteristics of the one or more reagents 40. For example, the one ormore reagents 40 may further comprise one or more components selectedfrom the group consisting of carrier proteins such as bovine serumalbumin (BSA), stabilizing agents, antimicrobial agents, a calcium salt,a potassium salt, a water soluble polymer, a sugar, gelatin, agarose, apolysaccharide, a saccharide, sucrose, polyethylene glycol, sodiumphosphate, glycine, an amino acid, antioxidants, a detergent, a buffersalt, and a buffer such as 4-(2-hydroxyethyl)-1-piperazineethanesulfonicacid (HEPES) buffer.

In accordance with different aspects of the present invention, the oneor more reagents 40 may include material suitable for inducing theintrinsic pathway. Materials suitable for inducing the intrinsic pathway(e.g., the aPTT or ACT analysis) may include one or more componentsselected from ellagic acid, celite, kaolin, diatomaceous earth, clay,silicon dioxide, synthetic or natural lipids, and synthetic or naturalphospholipids. In some embodiments a variety of other components may beincluded within the one or more reagents 40 to contribute tostabilization and/or deposition/dissolution characteristics of the oneor more reagents 40. For example, the one or more reagents 40 mayfurther comprise one or more components selected from the groupconsisting of dextran, dextrin, tergitol, buffers, a carrier protein, anamino acid, stabilizers, antimicrobials, antioxidants, a detergent, asaccharide, a polysaccharide, sucrose, polyethylene glycol, derivativesof polyethylene glycol, glycine, gelatin, buffer such as4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES) buffer,rhamnose, trehalose, and sugars.

In accordance with some aspects of the present invention, the one ormore substrates 45 used in the electrogenic assay may have an amidelinkage that mimics the thrombin-cleaved amide linkage in fibrinogen.Specifically, the one or more substrates 45 may comprise one or morethrombin-cleavable peptides such as those selected from the groupconsisting of H-D-Phe-Pip-Arg, H-D-Chg-Abu-Arg, CBZ-Gly-Pro-Arg,Boc-Val-Pro-Arg, H-D-Phe-Pro-Arg, Cyclohexylglycine-Ala-Arg,Tos-Gly-Pro-Arg, Bz-Phe-Val-Arg, Boc-Val-Pro-Arg, Ac-Val-Pro-Arg,Ac-Val-Hyp-Arg, Ac-(8-amino-3,6,dioxaoctanoyl-Val-Pro-Arg,Ac-Gly-Pro-Arg, Ac-(8-amino-3,6,dioxaoctanoyl-Gly-Pro-Arg,Ac-Gly-Hyp-Arg and H-D-Chg-Abu-Arg. Thrombin typically cleaves the amidebond at the carboxy-terminus of the arginine residue because the bondstructurally resembles the thrombin-cleaved amide linkage in fibrinogen.The product of the thrombin-substrate reaction includeselectrochemically inert compounds such as Tos-Gly-Pro-Arg,H-D-Phe-Pip-Arg, and/or Bz-Phe-Val-Arg- and electroactive compounds ordetectable moieties, preferably selected from the group consisting ofp-aminophenol, a quinone, a ferrocene, ferrocyanide derivative, otherorganometallic species, p-nitroaniline, o-dianisidine, 4,4′-bensidine,4-methoxy-2-naphthylamine, N-phenyl-p-phenylenediamine,N-[p-methoxyphenyl-]-p-phenylenediamine, and phenazine derivatives. Thetripeptide sequence was chosen because it renders the substratevirtually non-reactive with blood proteases other than thrombin and thereactivity of thrombin with the arginine amide linkage in the moleculeis very similar to its reactivity with the target amide linkage infibrinogen. When the one or more substrates 45 are present in a blood orblood derivative fluid sample or biological sample, generated activethrombin from activation of the coagulation pathway(s) via the one ormore reagents 40 simultaneously converts the one or more substrates 45and fibrinogen to their cleavage products. The electrochemical speciesreaction product is detected by the one or more transducers 37, e.g., anelectrochemical transducer.

Micro-Environment Sensor Structures

As discussed herein, micro-environment sensor structures comprise one ormore reagents and one or more substrates in any of a number of differentarrangements such that the introduction of the fluid sample, e.g., wholeblood, to the one or more reagents and the one or more substrates islocalized to the one or more sensors. In particular, themicro-environment sensor structures are configured to physicallyseparate the one or more reagents and/or reaction products from oneanother to avoid cross-activation of the cascade pathways or othercross-sensor interference once the one or more reagents have becomeexposed to the fluid sample.

As shown in FIG. 2, traditional POC coagulation assays have employed thereagent/substrate 60 printed as a dry substance on a wall 65 (e.g., acover) of a conduit that is opposite a surface of a sensor 70. The fluidsample 75 would need to be mixed with the dry substance, e.g., by pumposcillation, to dissolve the reagent/substrate 60 into the fluid sample75 and generate a mixture 80, which may be in the form of a gradientfrom a top of the conduit down to the sensor 70. However, such aconfiguration has at least three issues or disadvantages. Firstly, onlya small portion of the electroactive product generated via mixture 80will reach the surface of the sensor 70 and be oxidized, and thus amajority of the electroactive product will not be utilized. As a result,the usage of the reagent/substrate 60 is not efficient. Further, thefluid sample 75 is adulterated with the reagent/substrate 60, which maybe undesirable due to its possible impact with other sensors that maycome in contact with the fluid sample 75 (e.g., cross-sensorinterference). Secondly, in order to achieve adequate analyticalprecision, the reagent/substrate 60 should be dispersed uniformly in thefluid sample 75 as rapidly as possible. This may be a challenge forpoint-of-care devices where space and efficiency of mixing can belimited. It is especially true when the reagent/substrate 60 is in solidform and in a very small space relative to a volume of the fluid sample75. Thirdly, there is a possibility that the substrate interferes withthe reagent and/or coagulation factors. For example, mixing thesubstrate along with the reagent into the sample 75 before thecoagulation cascade has been initiated may manifest such interference.

In contrast to the traditional POC coagulation assays, some embodimentsof the present invention, as shown in FIG. 3, present the reagent 85associated with a substrate layer 90 formed in a localized manner nearthe surface of the sensor 95. For example, as shown in FIG. 3, thereagent 85 and the substrate 90 may be printed as a dry substancedirectly on a surface of the sensor 95. The fluid sample 100 may reactwith the reagent 85 and the substrate 90 without mixing (e.g., viapassive diffusion) (although some degree of mixing, e.g., fluidoscillation, may be desired), in a localized manner creating a gradientfrom the sensor 95 to a top of the conduit. Advantageously, thisarrangement of the reagent and the substrate presented directly on asurface of the sensor allows for a majority of the electroactive productto be oxidized, and thus utilized at the surface of the sensor. Thissensor arrangement is also beneficial due to the smaller sample volumerequired in the immediate sensor environment, and thus yielding a moreconcentrated reagent-to-sample assay zone.

Nonetheless, some of the issues (e.g., mitigation of cross-sensorinterference and substrate interference) apparent within the traditionalPOC coagulation assays may not be overcome by the arrangement shown inFIG. 3. For example, any reaction occurring in the immediate vicinity ofthe sensor could potentially interfere with the reagent and/orcoagulation factors and/or possibly with another reaction occurring atan adjacent sensor (i.e., a sensor within the same conduit and withinapproximately 3 mm of the sensor shown in FIG. 3). As such, this type ofsensor arrangement would not be characterized as a micro-environmentsensor. However, these remaining issues may be overcome via advancedmicro-fluidic systems of the present invention (e.g., splitting a singlesample into two or more parts and controlling movement of those partsinto two or more conduits or conduits), as discussed hereafter indetail, and/or appropriate spacing of sensors from one another. Forexample, in some embodiments, where adjacent sensors are covered by asame quiescent sample fluid, to prevent cross-sensor interference belowa given threshold, e.g., below 1%, it may be suitable to use modelsbased on a known diffusion coefficient for the interferent and theoverall assay time to determine an appropriate separation distancebetween sensors. In other embodiments, where the sample isnon-quiescent, other models for dynamic mixing may be suitable for useto select an appropriate sensor separation distance.

In additional or alternative embodiments, immobilizing the substrate 90on the sensor 95 has been unexpectedly demonstrated to address many orall of the above-mentioned issues. In accordance with these aspects ofthe present invention, the immobilization may be realized bycrosslinking (e.g., ultra-violet light, glutaraldehyde, etc.),entrapment, covalent binding, etc. One example of such amicro-environment arrangement is shown in FIG. 4 where the substrate 90is immobilized on the surface of the sensor 95 using a polymer layer105. In some embodiments, the immobilization may be performed by coatingthe sensor 95 with a polymer layer 105 that includes the substrate 90such that the substrate 90 is immobilized via the polymer layer 105 onthe surface of the sensor 95. In other words, the substrate 90 is formedas an immobilized porous substrate-polymer layer on the surface of thesensor 95 to create a vessel for maintaining the reaction of the fluidsample 100, the reagent 85, and the substrate 90 in a localized manneron a surface of the sensor 95. The fluid sample 100 may react with thereagent 85 and the substrate 90 without mixing (although some degree ofmixing, e.g., fluid oscillation, may be desired) in a localized mannerwithin the confines of (or above, and then diffused into) the polymerlayer 105 formed on the sensor.

Advantageously, this arrangement of the immobilized substrate presenteddirectly on a surface of the sensor allows for a majority of theelectroactive product to be oxidized, and thus utilized at the surfaceof the sensor. Even more advantageously, this arrangement of theimmobilized substrate provides for a micro-environment capable ofmaintaining the substrate and the electroactive product in the immediatevicinity of the sensor, and thus mitigating cross-sensor interferencewith an adjacent sensor during normal usage. Other potential benefits ofimmobilizing the substrate on the sensor include mitigation of substrateinterference via separation of the substrate from the reagent, reductionof material use, simplification of hardware and sensor design, andimprovement of product robustness.

FIGS. 5 and 6 provide empirical evidence that immobilizing the substratecan increase the response current and improve precision of analytedetection significantly. Specifically, FIG. 5 shows aPTT response curveswhere the x-axis is time/seconds and the y-axis is current/pA. In thisexample, the substrate was printed on a sensor, one immobilized with PVA(aPTT response curves 106) and the other not immobilized (aPTT responsecurves 107). An aPTT reagent was spiked into the whole blood. Aftermixing for about 30 seconds the sample was drawn from the sample tubeand filled into cartridges for testing. The electric current of theimmobilized substrate sensor (aPTT response curves 106) was over 30 nA,whereas that of the non-immobilized substrate sensor (aPTT responsecurves 107) was only about 3 nA. Their coefficient of variations of tMid(time at which the current reaches its middle point) were about 1% and2%, respectively. This data indicates that the presence of theimmobilized substrate directly over the sensor enabled the immediate andconcentrated redox reaction from the coagulation substrate leavinggroup. This in turn yielded faster clotting times and a more predictablesensor response.

Another example is shown in FIG. 6 where the x-axis is time/seconds andthe y-axis is current/pA. The PT response curves 108 represent theresponse of a non-immobilized substrate sensor to an i-STAT® PT controlfluid level 2, whereas the PT response curves 109 represent the responseof an immobilized substrate sensor to the same i-STAT® PT control fluidlevel 2. With respect to the non-immobilized substrate sensor, both thesubstrate and the reagent were printed together on the electrode, andmixed together with the sample during testing. With respect to theimmobilized substrate sensor, the substrate was immobilized with PVA onthe electrode and the reagent was printed on top of the immobilizedsubstrate, and there was no mixing during test. Use of the immobilizedsensor with a plasma control yielded a significant improvement inperformance such that the electric current increased from about 4 nA toabout 9 nA, and the coefficient of variation decreased from around 10%to about 3%. This data shows a significant improvement in the field ofcoagulation testing such that performance of a point of care devicecould now approach the performance of a central laboratory instrument(2-3% CV).

As shown in FIGS. 7A, 7B, and 7C, the micro-environment sensors of thepresent invention may have the reagent 110 and the immobilizedsubstrate-polymer layer 115 positioned in a number of differentarrangements with the components interacting with each other withoutmixing, although some degree of oscillation may be desired. For example,as shown in FIG. 7A, the reagent 110 may be positioned within orencapsulated by the immobilized substrate-polymer layer 115 (e.g., thereagent is integrated within the immobilized substrate-polymer layer).As shown in FIG. 7B, the reagent 110 may be coated over the immobilizedsubstrate-polymer layer 115 (e.g., the reagent is a separate layerdispensed on top of the immobilized substrate-polymer layer). As shownin FIG. 7C, the reagent 110 may be positioned substantially adjacent tothe immobilized substrate-polymer layer 115 and at least one transducerof the sensor 120 (e.g., the reagent is positioned within the conduitsuch that the reagent is abutted to or within an interactive distance ofthe substrate-polymer layer and/or the at least one transducer so as tostill function in conjunction with each other). As used herein, aninteractive distance means less than a longest dimension of the sensorwith the constraint of the reagent being positioned within a same planeor on a same wall/surface of a conduit as the sensor. Other variantswill also be apparent to those skilled in the art without departing fromthe spirit and scope of the present invention, for example the reagent110 may be formed as a combination of that shown in FIGS. 7B and 7C, oras shown in FIG. 7C with only part of the reagent 110 shown in FIG. 7B.

As shown in FIG. 8, in some embodiments, the present invention may bedirected to an analysis cartridge 125 comprising an inlet chamber 130configured to receive a fluid sample 135 and a conduit 140 fluidicallyconnected to the inlet chamber 130 and configured to receive the fluidsample 135 from the inlet chamber 130. The conduit 140 may comprise anarray of micro-environment sensors, e.g., a first micro-environmentsensor 145 and a second micro-environment sensor 150. The firstmicro-environment sensor 145 may comprise a first reagent 155 and afirst substrate 160 (e.g., a substrate immobilized within a polymerlayer) configured to detect a first diagnostic clotting time. Forexample, the first micro-environment sensor 145 may be a PT sensorcomprising a first reagent 155 that includes one or more components, asdiscussed herein, specific for triggering the extrinsic coagulationpathway and a first substrate layer 160 comprising a thrombin-cleavablepeptide with a detectable moiety as discussed herein. The secondmicro-environment sensor 150 may comprise a second reagent 165 and asecond substrate 170 (e.g., a substrate immobilized within a polymerlayer) configured to detect a second diagnostic clotting time. Forexample, the second micro-environment sensor 150 may be an aPTT sensorcomprising a second reagent 165 that includes one or more components, asdiscussed herein, specific for triggering the intrinsic coagulationpathway and a second substrate layer 170 comprising a thrombin-cleavablepeptide with a detectable moiety (e.g., a reagent and a substrateimmobilized within a polymer layer). As should be understood, althoughthe above-described analysis cartridge 125 is discussed with respect toa PT sensor and an aPTT sensor, various combinations and numbers ofsensors, e.g., a PT sensor, an aPTT sensor, and an ACT sensor, arecontemplated by the present invention without departing from the scopeof the present invention. For example, the first micro-environmentsensor 145 may be a PT sensor, and the second micro-environment sensor150 may be an aPTT sensor or an ACT sensor. In another aspect, the firstmicro-environment sensor 145 is an aPTT sensor, and the secondmicro-environment sensor 150 is a PT sensor or an ACT sensor. In anotheraspect, the first micro-environment sensor 145 is an ACT sensor, and thesecond micro-environment sensor 150 may be an aPTT sensor or a PTsensor. In still other embodiments, one of the micro-environment sensorsis a PT sensor, an aPTT sensor, or an ACT sensor, and another of thesensors is a sensor for detecting an analyte, related or unrelated tocoagulation.

Advantageously, the micro-environment sensor structures of the presentinvention are configured to physically separate the one or more reagentsand substrates to avoid cross-activation and/or interference of thecascade pathways once the one or more reagents and substrates havebecome exposed to the fluid sample. Even more advantageously,incorporation of the immobilized substrate and/or reagent polymer layerinto the coagulation assays provides for the ability to perform thecoagulation assays without requiring or while minimizing mixing, e.g.,oscillation of the sample fluid in a conduit, because coagulationactivation occurs in a localized and concentrated area over the sensorwith subsequent propagation of the test reaction into the immobilizedlayer, ultimately resulting in oxidation at the transducer.

Immobilized Substrate-Polymer Layer

In preferred embodiments, in order to physically separate the one ormore assays from one another to avoid cross-activation and promotelocalization of electrochemical or optical signals over the transducers,an immobilized substrate and/or reagent-polymer layer may be selectivelypatterned onto the sensors (e.g., coated over the transducer or workingelectrode/optical detector). As shown in FIG. 9, the immobilized polymerlayer 175 may be formed by either spin coating or by microdispensing.More specifically, an aqueous polymer matrix comprising one or morereagents and substrates and a polymer, such as a photoformable polymer(e.g., polyvinylalcohol (PVA)), may be utilized for immobilizing the oneor more substrates on or near the transducer 180. Additives includingbut not limited to a protein such as BSA, a sugar or sugar alcohol, suchas sucrose, sorbitol, or mannitol, may also be included in the aqueousmatrix. To those skilled in the art of polymer chemistry, the additionof some substances to the polymer layer(s) results in a number ofalterations to, including but not limited to, swelling reactions,diffusion coefficients, molecule stability, porosity, transport,reaction kinetics and the like. These alterations can be used tomodulate the micro-environment sensor response as required.

In accordance with some aspects of the invention, the one or moresubstrates may comprise one or more thrombin-cleavable peptides selectedfrom the group consisting of H-D-Phe-Pip-Arg, H-D-Chg-Abu-Arg,CBZ-Gly-Pro-Arg, Boc-Val-Pro-Arg, H-D-Phe-Pro-Arg,Cyclohexylglycine-Ala-Arg, Tos-Gly-Pro-Arg, Bz-Phe-Val-Arg,Boc-Val-Pro-Arg, Ac-Val-Pro-Arg, Ac-Val-Hyp-Arg,Ac-(8-amino-3,6,dioxaoctanoyl-Val-Pro-Arg, Ac-Gly-Pro-Arg,Ac-(8-amino-3,6,dioxaoctanoyl-Gly-Pro-Arg, Ac-Gly-Hyp-Arg andH-D-Chg-Abu-Arg. Optionally the two or more of these substrates may bemixed to obtain the thrombin activities and diffusional propertiesdesired in the immobilized substrate and/or reagent polymer layer.

In accordance with some aspects of the invention, the polymer thatcontains the substrate may comprise one or more materials, optionally inmatrix form. The material for the polymer, for example, may be selectedfrom the group consisting of PVA, styrylpyridinium polyvinylalcohol(SBQ-PVA), agarose, polyacrylamide, polymethyl methacrylate,N-methylpyrrolidone, polyvinylpyrrolidone, polyimide, a film-forminglatex, sepharose, polyurethanes, acrylates, methacrylates, polyethyleneglycols, polylactic acid, poly(lactic co-glycolic acid), hydroxypropylcellulose, celluloses, derivatives of cellulose,hydroxypropylmethylcellulose acetate succinate, inulin, fructans,derivatives of fructans, polyglycolic acid, Elvace, carboxymethylcellulose, polylactic acid, and poly(lactic co-glycolic acid). In someembodiments in which the material for the polymer comprises celluloses(e.g., hydroxypropyl cellulose), additives such as a plasticizer (e.g.,triethyl citrate, acetyl triethyl citrate, propylene glycol, glycerin,trimethylolpropane, polyethylene glycols, fatty acids, and derivativesthereof) and/or crosslinkers (e.g., carboxylic acids, glyoxal, and anyresin which is reactive with the available hydroxyl groups of thecellulose) may also be included in the aqueous matrix. Crosslinking ofthe materials may also affect the polymer layer swelling, permeability,diffusion, reaction kinetics etc. in order to modulate the sensorresponse as required.

Further to selection of the material for the polymer, another benefit ofimmobilizing the substrate and/or reagent includes using theimmobilizing matrix as a localized interferant neutralizer. For example,the selection of the material for the polymer may be dependent upon thetype of diagnostic clotting test to be performed using the immobilizedpolymer layer. For example, advantageously and unexpectedly it has beenfound that inclusion of cross-linked or non-cross-linked SBQ-PVA in theimmobilized polymer layer imparts a heparin neutralizing property orheparin insensitivity into the immobilized polymer layer. Consequently,in embodiments in which the diagnostic clotting test to be performedusing the immobilized polymer layer is a heparin sensitive test (e.g.,the PT test is known to be moderately sensitive to clot inhibitors suchas heparin), the polymer may be selected to be a heparin-neutralizingpolymer such as cross-linked or non-cross-linked SBQ-PVA. In someembodiments, the PVA may be a photo-activated stilbizonium salt.

FIGS. 10 and 11 exemplify this concept as follows. A coagulation sensorwas directly printed either with a large (FIG. 10) or small (FIG. 11)SBQ-PVA immobilized coagulation substrate. Subsequently, a PTcoagulation activator was printed on top of the immobilized substratematrix. These sensors were then tested with appropriate fluids to assesscoagulation time. As with the results from FIGS. 5 and 6, one benefit ofprinting more of the immobilized substrate matrix is that the increasedamount of immobilized substrate matrix led to an increase in pA. Also,when heparin is spiked into the whole blood sample, the clotting timesare not as extended as they are expected to be (FIGS. 10 and 11(response curves 181) (heparin spike) should be more similar to responsecurves 182 (abnormally long control fluid) clotting times, but areactually behaving more like whole blood without heparin, (responsecurves 183)). These results reflect the heparin neutralizing effect of apolymer such as SBQ-PVA. Further, applying a larger amount of theimmobilizing matrix results in a greater neutralizing effect (e.g.,compare the bias between whole blood and whole blood with heparin inFIGS. 10 and 11). FIG. 10 data represents a large matrix print and showsthat there is 14% extension when heparin is added to 1 IU/mL, while inFIG. 11 there is 55% extension when a smaller print amount of theneutralizing matrix is deposited. In both FIGS. 10 and 11, the heparinspiked blood behaves more like unaltered whole blood than it does likean abnormally long control fluid, reflecting that there is aninterferent/heparin neutralizing effect. This data shows that increasingthe area of the PVA print reduces the interferent effect of the heparinto the PT assay.

FIG. 12 provides further empirical evidence that a cross-linked SBQ-PVAmay be used to impart a heparin neutralizing property or heparininsensitivity to an immobilized substrate PT assay. Specifically, FIG.12 shows results of a PT assay performed on a whole blood sample(response curve 184) and a whole blood sample (response curve 185)spiked with 0.4 (response curve 184) or 1.2 IU/ml (response curve 185)of heparin without use of heparinase (a reagent conventionally includedin a PT test for neutralizing heparin), but with the addition ofincreasing amounts of SBQ-PVA to the same size print. The data showsthat increasing the concentration of the SBQ-PVA layer results in thelowering of the clotting time extensions from whole blood in thepresence of heparin. This shows that the PT micro-environment sensorresponse can be modulated to reduce interference of heparin without theuse of expensive heparinase.

Without being bound by theory, it appears that a positive chargeimparted by the cross-linked or non-cross-linked SBQ-PVA may impart theheparin neutralizing property or heparin insensitivity to theimmobilized substrate PT assay. More particularly, the SBQ pendent groupis a cation, the PVA is an anion, and the heparin is an anion, and thusit is hypothesized that repulsion in the localized area of the electrodevia the positively charged SBQ excludes heparin from the immobilizedsensor micro-environment or the positively charged SBQ interacts withthe negatively charged heparin and incapacitates the ability of heparinto act on coagulation factors. This theory is further evidenced by thefact that an anionic polymer such as hydroxypropyl cellulose may be usedfor diagnostic clotting time tests that monitor heparin therapy (e.g.,aPTT and ACT) without imparting a heparin neutralizing property orheparin insensitivity to the assay.

In additional or alternative embodiments, the polymer may be anon-heparin neutralizing polymer that could then be subsequently treatedor modified to become heparin neutralizing. For example, in embodimentsin which the diagnostic clotting test to be performed using theimmobilized polymer layer is heparin sensitive, the polymer may beselected to include at least one non-heparin neutralizing component, forexample, selected from the group consisting of hydroxypropyl cellulose,and Elvace, carboxymethyl cellulose, polylactic acid, polylactic acid,poly(lactic co-glycolic acid), celluloses, derivatives of cellulose,hydroxypropylmethylcellulose acetate succinate, inulin, fructose,fructans, derivatives of fructans, and polyglycolic acid. The one ormore components of the non-heparin neutralizing polymer may then betreated or modified to generate a heparin neutralizing layer. In someembodiments, the treatment or modification may include changing thecharge of the one or more components of the non-heparin neutralizingpolymer, adding heparinase to the polymer matrix, and/or configuring thepolymer layer to preferentially bind sulfate groups on the heparin.

In additional or alternative embodiments, the polymer may be formed of anon-heparin neutralizing polymer. For example, in embodiments in whichthe diagnostic clotting test to be performed using the immobilizedpolymer layer is for monitoring heparin therapy (e.g., the aPTT and ACTtest), the polymer layer may include at least one non-heparinneutralizing component optionally selected from the group consisting ofhydroxypropyl cellulose, Elvace, carboxymethyl cellulose, polylacticacid, polylactic acid, poly(lactic co-glycolic acid), celluloses,derivatives of cellulose, hydroxypropylmethylcellulose acetatesuccinate, inulin, fructans, derivatives of fructans, and polyglycolicacid.

In embodiments comprising spin coating, the immobilized substrate and/orreagent polymer layer may be photolithographically patterned usingultraviolet light to crosslink the material using a mask followed byremoval of the non-crosslinked material such that the immobilizedsubstrate and/or reagent polymer layer is selectively coated. Inembodiments comprising microdispensing (See, for example, U.S. Pat. No.5,554,339, which is incorporated herein by reference in its entirety),an appropriate quantity of each coating may be applied to an areaoptionally circumscribed by an additional structural componentconfigured as a containment boundary. Alternatively, surfacestreatments, e.g., exposure to gas plasmas, may be used to control thesurface energy, and thus the spreading of the microdispensed material.

In some embodiments, the one or more reagents and substrates 187 may beimmobilized within the polymer layer 175 as shown in FIG. 13A. Inaccordance with these aspects of the present invention, the aqueoussubstrate-polymer-reagent matrix comprising one or more substrates, apolymer such as photoformable polymer (e.g., PVA), and one or morereagents may be utilized for immobilizing the one or more substrates andthe one or more reagents on or near the transducer 180. The immobilizedpolymer layer 175 may be formed by either spin coating or bymicrodispensing the aqueous substrate-polymer-reagent matrix. Inpreferred embodiments, the one or more reagents or substrates 187 for anaPTT or ACT test may be immobilized within the polymer layer 175 and thedried volume of the immobilized reagent-substrate-polymer layer 175comprising the one or more reagents or substrates 187 may be in therange of about 0.55-2.0 nL, preferably in the range of about 1.0-1.5 nL.In some embodiments, the immobilized polymer layer 175 is substantiallyplanar and has a thickness in the range of about 0.1-100 μm. Inadditional or alternative embodiments, the immobilized polymer layer 175is substantially domed and has a maximum thickness of the dome in therange of about 0.1-100 μm. Although the reagents are shown in FIG. 13Aheterogeneously localized in the central region of the polymer layer175, in preferable embodiments the reagent(s) is homogeneously dispersedthroughout the substrate-polymer layer.

In some embodiments, the one or more reagents or substrates 187 may beformed as a separate layer over and/or adjacent to the immobilizedpolymer layer 175 as shown in FIGS. 13B and 13C. Further, the one ormore regents or substrates may be localized/immobilized together or inseparate locations. In accordance with these aspects of the presentinvention, the one or more reagents or substrates 187 may be spin coatedor printed over and/or adjacent to the immobilized polymer layer 175(e.g., the PVA layer) to localize electrochemical or optical signalsover or near the transducer 180. In preferred embodiments, the one ormore reagents or substrates 187 for a PT test may be formed separatefrom the immobilized polymer layer 175 and the dried volume of theimmobilized polymer layer 175 may be in the range of 1.5-2.2 nL,preferably in the range of 1.60-2.00 nL. In some embodiments, theimmobilized polymer layer 175 is substantially planar and has athickness in the range of about 0.1-100 μm. In additional or alternativeembodiments, the immobilized polymer layer 175 is substantially domedand has a maximum thickness of the dome in the range of about 0.1-100μm.

Sensor and Chip Design

A preferred embodiment of a microfabricated sensor array comprises atleast one transducer (e.g., a working electrode or optical detector).For example, the microfabricated sensor array may comprise a pair ofmicro-environment sensors or transducers comprising a firstmicro-environment sensor or transducer (e.g., a PT sensor) andoptionally a second micro-environment sensor or transducer (e.g., anaPTT sensor). In some embodiments, the micro-environment sensors ortransducers may be fabricated as adjacent structures, respectively, on asilicon chip.

In additional or alternative embodiments, the microfabricated sensorarray may further comprise in addition to the first micro-environmentsensor or transducer and optionally the second micro-environment sensoror transducer, one or more blood chemistry sensors. For example, thesensor array may further comprise one or more of sensors configured tomeasure one or more of sodium, potassium, calcium, chloride, carbondioxide, glucose, blood urea nitrogen (BUN), creatinine, pH, partialpressure CO₂, partial pressure O₂, lactate, magnesium, or anotheranalyte.

In some embodiments, the transducers may be formed as electrodes withgold surfaces coated with a photo defined polyimide layer. For example,wafer-level micro-fabrication of a preferred embodiment of the sensorarray may be achieved as shown in FIG. 14. A planar non-conductingsubstrate 190 may be used as a base for the sensor array. A conductinglayer 195 may be deposited on the substrate 190 by conventional means,e.g., conductive printing, or micro-fabrication technique known to thoseof skill in the art to form at least one transistor. The conductinglayer 195 may comprise a noble metal such as gold, platinum, silver,palladium, iridium, or alloys thereof, although other unreactive metalssuch as titanium and tungsten or alloys thereof may also be used, asmany non-metallic electrodes of graphite, conductive polymer, or othermaterials may also be used.

For example, a base electrode may comprise a square array of 5-10 μmgold disks, e.g., 7 μm gold disks, on 15 μm centers. The array may covera region, e.g., a circular region, approximately 300 to 900 μm indiameter, optionally 400-800 μm or about 600 μm in diameter, and may beformed by photo-patterning a thin layer of polyimide or photoresist ofthickness up to 1.5 μm over a substrate made from a series of layerscomprising Si, SiO₂, TiW, and/or Au, or combinations thereof. In someembodiments, the base electrode has a working area of about 130,000 to300,000 sq μm, the volume of sample directly over the sensor may beabout 0.1-0.3 μL, and the volume of the sample over the chip may be 1-3μL. In accordance with these aspects of the present invention, theconduit in a region of the base electrode has a volume to sensor arearatio of less than about 6 μL to about 1 square mm, preferably less thanabout 50 mm to about 2 square mm, more preferably less than about 100 μmto about 500 square μm. Accordingly, the array of microelectrodesaffords high collection efficiency of a detectable moiety that is anelectroactive species with a reduced contribution from anyelectrochemical background current associated with the capacitance ofthe exposed metal. In particular, openings in the insulating polyimideor photoresist layer define a region of gold electrodes at which theelectroactive species, e.g., p-aminophenol, may be oxidized such as in atwo electron per molecule reaction.

Micro-fabrication techniques (e.g. photolithography and plasmadeposition) may be utilized for construction of the multilayered sensorstructures in confined spaces. For example, methods formicro-fabrication of electrochemical immunosensors on silicon substratesare disclosed in U.S. Pat. No. 5,200,051, which is hereby incorporatedby reference in its entirety, and include, for example, dispensingmethods, methods for attaching substrates and reagents to surfacesincluding photoformed layers, and methods for performing electrochemicalassays.

The microfabricated sensor array may also comprise an electricalconnection 195 and an immobilized polymer layer 205 (as discussed abovewith respect to FIGS. 4, 7A, 7B, and 7C), which is deposited onto atleast a portion of the conducting layer 195 and/or the non-conductingsubstrate 190. In the present invention, the immobilized polymer layer205 may be a porous polymer layer comprising a thrombin-cleavablepeptide with a detectable moiety that is configured to respond to thepresence of active thrombin by producing a change that is capable ofbeing measured.

As shown in FIGS. 15 and 16, in some embodiments, the microfabricatedsensor array may comprise a silicon chip 210 that includesmicro-environment amperometric sensors or transducers 215 and 220located on different vertical planes (a) and (b) of the silicon chip210. The sensor 215 may be connected via wiring 225 to a firstamperometric pin 230 (e.g., temporary electrical connector) and thesensor 220 may be connected via wiring 235 to a second amperometric pin240 (e.g., temporary electrical connector). In some embodiments, thesensor 215 may be configured as an aPTT sensor and the sensor 220 may beconfigured as a PT sensor both of which are formed on the single siliconchip 210 and positioned within one or more conduits of the point of caretest cartridge. As illustrated in FIG. 15, the sensor 215 may beconstructed with a target reticle design preferably comprising aplurality of concentric rings (e.g., 2, 3, 4 or more concentric rings)in an upper area of the silicon chip 210 and the sensor 220 may beconstructed with a target reticle design preferably comprising aplurality of concentric rings (e.g., 2, 3, 4 or more concentric rings)in a lower area of the silicon chip 210. Specifically, the design andarrangement of the sensors 215 and 220 on the chip 210 are selectedbased on printing and performance characteristics for each of thesensors 215 and 220. However, it should be understood to those ofordinary skill in the art that any design or arrangement for the sensorsis contemplated without departing from the spirit and scope of thepresent invention. Furthermore, although the sensors 215 and 220 in theexample in FIG. 15 are amperometric sensors, other electrochemicalprocesses or optical processes which use other electrochemical oroptical sensors, e.g., optical wave guides and charge-coupled device(CCD) camera chips, can be used. For example, a potentiometric sensormay be used to detect ion species such as Na⁺ or K⁺.

As described herein, the amperometric sensors or transducers 215 and 220may be formed as electrodes with gold surfaces that are exposed (e.g.,no polyimide or photoresist covering) to the inside environment of theconduit and configured to directly contact a biological sample disposedwithin the conduit. The wirings 225 and 235 may be formed with goldsurfaces that are coated with a photo defined polyimide or photoresistlayer such that the wirings 225 and 235 are insulated from exposure tothe biological sample disposed within the conduit. The wirings 225 and235 may be formed comprising containment ring structures 245 and 250configured to contain the immobilized reagent-substrate-polymer layer.For example, the immobilized reagent-substrate-polymer layer (asdiscussed above with respect to FIGS. 4, 7A, 7B, and 7C) may bedeposited onto at least a portion of the sensors 215 and/or 220 withinthe containment ring structures 245 and/or 250. The wirings 225 and 235terminate at the first amperometric pin 230 and the second amperometricpin 240 respectively, which are used to make contact with a connector inan analyzer or cartridge reader (e.g., an i-STAT® cartridge reader asdescribed in U.S. Pat. No. 4,954,087, the entirety of which isincorporated herein by reference).

In the preferred embodiments of the present invention, the analyzerapplies a potential via the first amperometric pin 230 and the secondamperometric pin 240 between each of the amperometric sensors 215 and220 and a reference electrode (described in detail below with respect toFIG. 17), and measures current changes generated by cleaved substrate asan electrochemical signal. The electrochemical signal being proportionalto the concentration of the product in the biological sample. Theamperometric sensors 215 and 220 have an applied potential ofapproximately +0.4 V versus the reference electrode and, in anotherpreferred embodiment, the amperometric sensors 215 and 220 have anapplied potential of approximately +0.1 V versus the referenceelectrode. The signal generated by the enzyme reaction product atapproximately +0.1V is distinguishable from the signal generated by theunreacted substrate at approximately +0.4 V.

In the embodiments of the invention which use the thrombin cleavablepeptide Tos-Gly-Pro-Arg-, H-D-Phe-Pip-Arg, or Bz-Phe-Val-Arg attached toan N-phenyl-p-phenylenediamine orN-[p-methoxyphenyl-]-p-phenylenediamine detectable moiety, the intactsubstrates are detected at a voltage of approximately +0.4V. Theelectrogenic reaction products N-phenyl-p-phenylenediamine orN-[p-methoxyphenyl-]-p-phenylenediamine are detected at a voltage ofapproximately +0.1V. Thus in these embodiments, the analyzer applies apotential to the amperometric sensors 215 and 220 with the generation ofan electrochemical signal which is proportional to the concentration ofthe substrate in the biological sample. Also, the analyzer applies apotential to the amperometric sensors 215 and 220 with the generation ofan electrochemical signal which is proportional to the concentration ofthe product in the biological sample. After hydrolysis of the substrateby thrombin, a product is formed which reacts at the amperometricsensors 215 and 220 with the generation of a signal distinguishable fromthe signal generated by the substrate.

It should be noted that the exact voltages used to amperometricallydetect the substrate and the product will vary depending on the chemicalstructure of the substrate and product. It is important that thedifference in the voltages used to detect the substrate and the productbe great enough to prevent interference between the readings. With somesubstrates, the voltage required to electrochemically detect thesubstrate is so high as to be beyond practical measurement in an aqueousbuffered solution. In these cases, it is only necessary that the productbe detectable amperometrically.

In some embodiments, the silicon chip 210 shown in FIG. 15 may furtherinclude multi-conduit conductometric sensors 255 and 260 (e.g.,hematocrit sensors). The conductimetric sensors 255 and 260 areconfigured to determine biological sample arrival and/or departure atthe amperometric sensors 215 and 220. More specifically, theconductometric sensors 255 and 260 lie perpendicular to a length of theconduit or sensor conduit, and an electrical resistance between pairs ofelectrodes for each sensor may be used to monitor a relative position ofa fluid front of the biological sample. At the extremes, an open circuitreading indicates that the biological sample has been pushed off theamperometric sensors 215 and 220 and a closed circuit reading indicatesthe amperometric sensors 215 and 220 are covered with the biologicalsample.

As shown in FIG. 15, the conductometric sensor 255 may comprise at leasttwo electrodes 265 and 270 (i.e., first electrode pair) positionedupstream of a midpoint of the amperometric sensor 215. The electrodes265 and 270 may be connected via wirings 275 and 280 to a conductometriclow pin 285 and an AC source or conductometric high pin 290,respectively (e.g., temporary electrical connectors). The wirings 275and 280 may be formed with a gold surface that is coated with a photodefined polyimide or photoresist layer such that the wirings 275 and 280are insulated from exposure to the biological sample disposed within theconduits. The conductometric sensor 260 may comprise at least twoelectrodes 295 and 300 (i.e., second electrode pair) positioneddownstream of a midpoint of the amperometric sensor 220. The electrodes295 and 300 may be connected via wirings 275 and 280 to a conductometriclow pin 285 and an AC source or conductometric high pin 290,respectively (e.g., temporary electrical connectors). As such, in someembodiments, the fluid reaches the first electrode pair in a firstfluidic conduit (e.g., prior to arriving at amperometric sensor 215),then subsequently arrives at the second electrode pair in a secondfluidic conduit (e.g., after arriving at amperometric sensor 220).

As shown in FIG. 16, in another embodiment, the silicon chip 210 mayfurther include a third conductometric sensor 301 comprising at leasttwo electrodes 302 and 303. The electrodes 302 and 303 may be connectedvia wiring 304 to a second AC source or conductometric high pin 305(e.g., temporary electrical connector). In accordance with these aspectsof the present invention, the use of a third sensor allows for twobinary fluid detection events, e.g., both are OFF/ON, which is easilydetectable with the current circuitry and software limitations. In thecase of two conductivity sensors (shown in FIG. 15), the currentcircuitry and software relies on the ability to detect two ‘drops’ inthe resistance of the sample in quick succession. Typically the firstdrop is large as it goes from a dry state to a wet state and the circuitis completed. The second drop in resistance, when the sample arrives inthe second fluidic conduit, is much smaller and therefore more difficultto differentiate from signal noise and small changes in the signal. Inaddition, the amplitude of each resistance change varies depending onthe sample properties. Accordingly and advantageously, in someembodiments, the arrangement of having three conductometric sensorsallows for two switchable conductivity paths using the conductometricsensor 255 (shown in FIG. 15) and the conductometric sensor 301 (shownin FIG. 16).

As shown in FIG. 17, in some embodiments, the microfabricated sensorarray may further comprise a ground chip 306 that includes a referencesensor or electrode 307. In accordance with aspects of the presentinvention, in which the sensors 215 and 220 are amperometric sensors,the reference electrode 307 may be configured as a counter electrode tocomplete the circuitry. In a preferred embodiment, the referenceelectrode 307 may comprise silver metal (Ag) and its silver salt (AgCl)deposited on a solid substrate (i.e., an Ag/AgCl reference electrode).The reference electrode 307 may be connected via wiring 308 to areference pin 309 (e.g., temporary electrical connector). Themicrofabricated sensor array may be designed such that the ground chip306 is positioned upstream of the semiconductor chip 210 as discussed infurther detail with respect to FIGS. 15 and 16. However, it should beunderstood that other arrangements for sensor and ground chips arepossible without departing from the spirit and scope of the presentinvention. For example, the sensor array may further comprise one ormore additional sensor chips (not shown) configured to detect variousanalytes of potential interest, such as troponin I, troponin T, CKMB,procalcitonin, bHCG, HCG, NTproBNP, proBNP, BNP, myoglobin, parathyroidhormone, d-dimer, NGAL, galectin-3, and/or PSA, among other analytes.

As shown in FIG. 18, in preferred embodiments, the microfabricatedsensor array may comprise a silicon chip 310 that includesmicro-environment amperometric sensors or transducers 315 and 320located on a same vertical plane (a) of the silicon chip 310. The sensor315 may be connected via wiring 325 to a first amperometric pin 330(e.g., temporary electrical connector) and the sensor 320 may beconnected via wiring 335 to a second amperometric pin 340 (e.g.,temporary electrical connector). In some embodiments, the sensor 315 maybe configured as an aPTT sensor and the sensor 320 may be configured asa PT sensor both of which are formed on a single chip 310 and positionedwithin the conduit of the point of care test cartridge. As illustratedin FIG. 18, the sensor 315 may be constructed with a donut shaped designin an upstream position to that of the sensor 320 constructed with atarget reticle design comprising a plurality of concentric rings (e.g.,2, 3, 4 or more concentric rings). Specifically, the design andarrangement of the sensors 315 and 320 on the chip 310 are selectedbased on printing and performance characteristics for each of thesensors 315 and 320. However, it should be understood to those ofordinary skill in the art that any design or arrangement for the sensorsis contemplated without departing from the spirit and scope of thepresent invention. Furthermore, although the sensors 315 and 320 in theexample in FIG. 18 are amperometric sensors, other electrochemicalprocesses or optical processes which use other electrochemical oroptical sensors can be used. For example, a potentiometric sensor may beused to detect ion species such as Na⁺ or K⁺.

As described herein, the sensors or transducers 315 and 320 may beformed as electrodes with gold surfaces that are exposed (e.g., nopolyimide or photoresist covering) to the inside environment of theconduit and configured to directly contact a biological sample disposedwithin the conduit. The wirings 325 and 335 may be formed with goldsurfaces that are coated with a photo defined polyimide layer such thatthe wirings 325 and 335 are insulated from exposure to the biologicalsample disposed within the conduit. The wirings 325 and 335 may beformed comprising containment ring structures 345 and 350 configured tocontain the immobilized reagent-substrate-polymer layer. For example,the immobilized reagent-substrate-polymer layer (as discussed above withrespect to FIGS. 4, 7A, 7B, and 7C) may be deposited onto at least aportion of the sensors 315 and/or 320 within the containment ringstructures 345 and/or 350. The wirings 325 and 335 terminate at thefirst amperometric pin 330 and the second amperometric pin 340respectively, which are used to make contact with a connector in ananalyzer or cartridge reader (e.g., an i-STAT® cartridge reader asdescribed in U.S. Pat. No. 4,954,087).

In some embodiments, the silicon chip 310 further includes an integratedreference electrode 355. In accordance with aspects of the presentinvention, in which the sensors 315 and 320 are amperometric sensors,the reference electrode 355 is configured as a counter electrode tocomplete the circuitry. The reference electrode 355 may comprise silvermetal (Ag) and its silver salt (AgCl) deposited on a solid substrate(i.e., a Ag/AgCl reference electrode). The reference electrode may beconnected via wiring 360 to an AC ground and reference pin 365 (e.g.,temporary electrical connector). The wiring 360 may be formed with agold surface that is coated with a photo defined polyimide orphotoresist layer such that the wiring 360 is insulated from exposure tothe biological sample disposed within the conduit. In preferredembodiments, the reference electrode 355 is designed in a checkerboardpattern as illustrated in FIG. 18 to improve wettability of a surface ofthe reference electrode 355. Specifically, it has been foundunexpectedly that the wettability of the reference electrode 355 may beimproved using the checkerboard pattern because AgCl is relativelyhydrophobic and can promote the formation of an air bubble over thesurface of the reference electrode 355 when a solid patch of AgCl isused, which results in a poor circuit.

As discussed in detail above with respect to the silicon chip 310 and asshown in FIG. 19, in the preferred embodiments of the present inventionthe analyzer applies a potential via the first amperometric pin 330 andthe second amperometric pin 340 between each of the amperometric sensors315 and 320 and the reference electrode 355, and measures currentchanges generated by cleaved substrate as an electrochemical signal. Theelectrochemical signal being proportional to the concentration of theproduct in the biological sample. The amperometric sensors 315 and 320have an applied potential of approximately +0.4 V versus the referenceelectrode 355 and, in another preferred embodiment, the amperometricsensors 315 and 320 have an applied potential of approximately +0.1 Vversus the reference electrode 355. The signal generated by the enzymereaction product at approximately +0.1V is distinguishable from thesignal generated by the unreacted substrate at approximately +0.4 V.

With reference back to FIG. 18, in some embodiments, the silicon chip310 may further include conductometric sensors 370 and 375 (which canalso function as hematocrit sensors). The conductimetric sensors 370 and375 may be split to form two sensor pairs with one at each end of thechip 310. The conductimetric sensors 370 and 375 are configured todetermine biological sample arrival and/or departure at the amperometricsensors 315 and 320, respectively. More specifically, the conductometricsensors 370 and 375 lie in an arc that is perpendicular to a length ofthe conduit or sensor conduit, and an electrical resistance betweenpairs of electrodes for each sensor may be used to monitor a relativeposition of a fluid front of the biological sample. At the extremes, anopen circuit reading indicates that the biological sample has beenpushed off the amperometric sensors 315 and 320 and a closed circuitreading indicates the amperometric sensors 315 and 320 are covered withthe biological sample.

As shown in FIG. 20, the conductometric sensor 370 may comprise at leasttwo electrodes 380 and 385 (i.e., first electrode pair) positioned at apredetermined distance (d1) from one another. In some embodiments, theconductometric sensor 370 may be positioned on the silicon chip 310relative to a midpoint (v) of the amperometric sensor 315 (e.g.,upstream, downstream, or in-line with the midpoint (v)). The electrode380 may be connected via wiring 390 to an AC source pin 395 (e.g.,temporary electrical connector). The electrode 385 may be connected viawiring 400, the reference electrode 355, and the wiring 360 to the ACground and reference pin 365. The wirings 390 and 400 may be formed witha gold surface that is coated with a photo defined polyimide orphotoresist layer such that the wirings 390 and 400 are insulated fromexposure to the biological sample disposed within the conduit.

The conductometric sensor 375 may comprise at least two electrodes 405and 410 (i.e., second electrode pair) positioned at a predetermineddistance (d2) from one another. In some embodiments, the conductometricsensor 375 may be positioned on the silicon chip 310 relative to amidpoint (x) of the amperometric sensor 320 (e.g., upstream, downstream,or in-line with the midpoint (x)). The electrode 405 may be connectedvia wiring 415, the reference electrode 355, and the wiring 360 to theAC ground and reference pin 365. The electrode 410 may be connected viawiring 420 and the wiring 390 to the AC source pin 395. The wirings 415and 420 may be formed with a gold surface that is coated with a photodefined polyimide or photoresist layer such that the wirings 415 and 420are insulated from exposure to the biological sample disposed within theconduit.

In preferred embodiments, the conductometric sensors 370 and 375 areconfigured to detect arrival of the biological sample within the conduitat the amperometric sensors 315 and 320, respectively. As shown in FIG.21, the arrival of the biological sample at the amperometric sensors 315and 320 may be detected based on determination of a first resistancedrop 425 when the biological sample reaches conductivity sensor 370 anda second resistance drop 430 when the biological sample reachesconductivity sensor 375. In additional or alternative embodiments,determination of a rise or spike (not shown) in the resistance at eitheror both of the conductometric sensors 370 and 375 may be used to detectthe presence of an air bubble within the conduit that is positioned overeither or both of the amperometric sensors 315 and 320.

A resistance profile for the conductometric sensors 370 and 375 shouldpreferably provide two well defined resistance drops of roughly equalamplitude. In some chip designs, as shown in FIG. 22A, conductometricsensors 435 and 440 may be configured as separate bars on opposite endsof the chip near respective amperometric sensors 445 and 450. However,the resistance profile 455 for such a design is found to often includean additional step 460, which is attributable to the sample temporarilystopping on the reference electrode 465 due to the hydrophobic nature ofthe reference electrode 465. As should be understood, this could make itdifficult to decipher the second resistance drop as either the wettingof the reference electrode 465 or the sample arriving at the secondconductometric sensor 440. Additionally, the time between the two stepsis quite short, making the timing difficult, and the resistance drop ofthe second arrival is much smaller compared to the first drop, makingthe detection difficult.

Accordingly, as shown in FIG. 22B, the chip design implemented inpreferred embodiments of the present invention utilizes theconductometric sensors 370 and 375, which are each split to comprise atleast two electrodes 380, 385 and 405, 410 spaced apart at predetermineddistances (d1) and (d2), respectively. As illustrated in the resistanceprofile 470, the dominant resistance drops 425 and 430 occur at the twopairs of conductometric sensors 370 and 375. Thus, reducing the impactof the additional resistance drop 460 (shown in FIG. 22A) observed fromthe wetting of the reference electrode 355. Further, the conductometricsensors 370 and 375 are placed at a front and rear of the chip toincrease a time between the resistance drops 425 and 430 to betterdifferentiate the resistance drops 425 and 430. Moreover, in someembodiments, the spacing or predetermined distance (d1) provided betweenthe electrodes 380 and 385 is a value “n” greater than that of thespacing or predetermined distance (d2) provided between the electrodes405 and 410 such that an amplitude of the second resistance drop 430 isincreased over a resistance drop 475 (shown in FIG. 22A) of thealternative chip design. For example, (d1) may be constructed twice aslarge as that of (d2) to achieve about a 1000 ohm increase in amplitudeof the second resistance drop. The increase in (d1) over that of (d2)effectively increases the ratio of resistance drops for the chip designshown in FIG. 22B over that of the ratio of resistance drops for thechip design shown in FIG. 22A. Advantageously, this increase inresistance drops allows for better detection of the arrival of thebiological sample at the conductometric sensors 370 and 375 during anon/forward motor or pump position 480.

In some embodiments, processes of the present invention may includecontinually moving the biological sample forward and back over the chipat a controlled velocity. Controlling the time for which theconductometric sensors 370 and 375 remain as open and closed circuitscontrols the position at which the biological sample changes direction.For example, a pneumatic pump within the analyzer may be configured tooscillate the biological sample in the conduit with the trailing edge ofthe biological sample positioned in the region of the conductometricsensor 370 in order to dissolve the substrate in that portion of thesample near the trailing edge. The oscillation may be at a frequency inthe range of 0.2 to 10 Hertz for a period in the range of 1 to 100seconds. In a preferred method, the oscillation may be at a frequency inthe range of about 1.5 Hertz for a period of about 20 seconds. Inanother preferred method the oscillation may be at a frequency of about0.3 Hertz and the amperometric sensors 315 and 320 (as shown in FIG. 20)may be configured to generate a signal at each oscillation. Iferythrocytes are present in the biological sample, the oscillation maybe at a frequency adequate to prevent the settling of erythrocytes onthe amperometric sensors 315 and 320.

In some embodiments, the amperometric sensors 315 and 320 determine theconcentration of product each time the biological sample is oscillatedpast the amperometric sensors 315 and 320. For example, a firstamperometric sensor signal may be stored by the analyzer for each of theamperometric sensors 315 and 320 and subsequent signals from theamperometric sensors 315 and 320 may be stored and compared to the firstand other stored signals in order to determine a maximum rate of changein the amperometric sensor signals. These data points may then beanalyzed to determine a fixed fraction of a maximum rate of change ofthe amperometric sensor signals. These data points may thus be used todetermine a coagulation parameter of interest for each of theamperometric sensors 315 and 320.

In alternative embodiments, the sensors or transducers may be formed asan optical detector, e.g. CCD camera chip and optical wave guide. Theoptical detector may either be a detector of fluorescence,chemiluminescence, or bioluminescence emission from the detectablemoiety or a detector of absorbance by the detectable moiety. In suchembodiments, the detectable moiety may be an optical dye, a fluorescenceemitter, chemiluminescence emitter or a bioluminescence emitter.

In other embodiments, the sensor or transducers may be formed as a teststrip, e.g., a glucose test strip, as described in U.S. patentapplication Ser. No. 13/724,348, which is incorporated herein in itsentirety. For example, a test strip may be included within thecartridges described herein. In some embodiments, the sample may bemanually placed on the test strip and, as such, the microfluidic systemsdescribed herein would not need to be included with such embodiments. Asis well known in the art, glucose test strip devices can include passivecapillary fluidic elements to deliver the sample to a sensor or sensorarray. As such, the elements, features, and functionality of a glucosetest strip could be adapted to the present invention without departingfrom the spirit and scope of the present invention.

Systems and Processes for Sample Analysis

As shown in FIG. 23, the system 500 of the present invention maycomprise a self-contained disposable sensing device or cartridge 505 anda reader device or instrument 510 (e.g., an analyzer). In someembodiments, the cartridge 505 is a single-use device configured to bedisposable after the single-use. A fluid sample (e.g., whole blood) tobe measured is drawn into a sample entry orifice or port 515 in thecartridge 505, and the cartridge 505 may be inserted into the reader 510through a slotted opening 520. The reader 510 may comprise a processorconfigured to perform measurements of analyte concentrations,measurements of resistances, identify analytes or sets of analytes thata chip is configured to measure, and/or determinations of diagnosticclotting time within the fluid sample, as discussed herein in furtherdetail. Measurements and determinations performed by the reader 510 maybe output to a display 525 or other output device, such as a printer ordata management system 530 via a port 535 on the reader 510 to acomputer port 540. Transmission can be via Wifi, Bluetooth link,infrared and the like. In embodiments where the sensors 545 in thecartridge 505, e.g., micro-environment sensors, are based onelectrochemical principles of operation, (e.g., a first sensor andoptionally a second sensor) may be configured to make electrical contactwith the reader 510 via an electrical connector 550. For example, theconnector may be of the design disclosed in jointly owned U.S. Pat. No.4,954,087, incorporated herein by reference in its entirety. In someembodiments, the PT and aPTT sensors may be configured to connect withan electrical connector of a test meter within the reader 510 via theelectrical connector 550 (see, e.g., U.S. Pat. Nos. 5,096,669 and4,954,087, incorporated herein by reference in their entireties). Thereader 510 may also include a method for automatic fluid flowcompensation in the cartridge 505, as disclosed in jointly owned U.S.Pat. No. 5,821,399, which also is incorporated herein by reference inits entirety.

In one embodiment, as shown in FIG. 24, the self-contained disposablesensing device or cartridge 555 may comprise a cover 560, a base 565,and a thin-film adhesive gasket (not shown) that is disposed between thebase 565 and the cover 560. The cartridge 555 may be configured forinsertion into the reader 510, and therefore the cartridge 555 maycomprise a plurality of mechanical and electrical connections (notshown) for this purpose. Advantageously, a feature of the cartridge 555is that once a fluid or biological sample is loaded within the cartridge555, analysis of the fluid or biological sample may be completed and thecartridge 555 may be discarded without an operator or others contactingthe fluid or biological sample.

Referring to FIG. 24, the cover 560 may be made of a rigid material,preferably plastic, and capable of repetitive deformation at flexiblehinge regions 570, 575, and 580 without cracking. The cover 560 maycomprise a lid 585, attached to a main body of the cover 560 by theflexible hinge 570. In operation, after introduction of the fluid orbiological sample into a sample holding chamber 590 through a sampleentry port 595, the lid 585 may be secured over an entrance to thesample entry port 595, preventing sample leakage. The lid 585 may beheld in place by a hook 600. The cover 560 may further comprise twodeformable members 605 and 610 that are moveable relative to the body ofthe cover 560, and which may be attached to the cover 560 by theflexible hinge regions 575 and 580.

The deformable member 610 may be configured to be operated by a firstpumping means such that a force is exerted upon an air bladder comprisedof cavity 615 and the gasket. Operation of the deformable member 610displaces fluid within conduits of the cartridge 555. The deformablemember 605 may be configured to be operated upon by a second pumpingmeans such that a force is exerted upon the gasket, which can deformbecause of slits cut therein. In some embodiments, deformation of thegasket may transmit pressure onto a fluid-containing foil pack filledwith a fluid, e.g., approximately 130 μL of analysis/wash solution, acontrol fluid, or calibrant fluid, located in cavity 620, rupturing thefoil pack, and expelling fluid into conduit 625 for subsequent use inother conduits during sample analysis. As should be understood, whilecoagulation assay formats do not generally require the use of thesefluids, the fluids may generally be required in a single device thatcombines coagulation tests with other tests, e.g., a wash fluid inimmunoassays for analytes such as BNP and troponin, and a calibrantfluid in chemistry tests such as potassium, creatinine and glucose. Inalternative embodiments, the deformation of the gasket may transmitpressure onto an air bladder comprised of cavity 620 for displacingfluid within conduits of the cartridge 555. In yet additionalembodiments, the second pumping means may not operate upon the cavity620, and instead, the cavity 620 may be configured as a waste chamber.

Additional action in the cartridge 555 generated by mechanisms withinthe reader 510 (discussed with respect to FIG. 23) applied to thecartridge 555 may be used to inject one or more air segments into thefluid or biological sample at controlled positions within the sampleholding chamber 590 and conduit 630. The air segments may be used towash a sensor surface of the sensor array and the surrounding conduit630 with a minimum amount of fluid (e.g., a limited wash cycle in whichthe volume of wash may be less than fifty times a volume of the fluid orbiological sample and/or fewer than three independent cycles of cleanwash buffer (e.g., three independent washing steps with fresh washbuffer)), as should be understood by those of ordinary skill in the artof immunoassay procedures. For example, the cover 560 may furthercomprise a hole covered by a thin pliable film. In operation, pressureexerted upon the film may expel one or more air segments into theconduit 630 through a small hole in the gasket. In some embodiments, across-sectional area of the conduit 630 may be in the range of about 0.1mm² to about 10 mm².

In some embodiments, a lower surface of the cover 560 further comprisessample holding chamber 590, the conduit 630 and another conduit 635(e.g., a waste conduit). The sample holding chamber 590 and the conduit630 may include one or more constrictions or capillary stops 640 and 642that control fluid flow by providing resistance to the flow of the fluidor biological sample. Optional coatings (not shown), e.g., dry reagentcoatings, may provide hydrophobic surfaces on the conduit 630, whichtogether with gasket holes control fluid flow between the sample holdingchamber 590 and the conduit 635. The sample holding chamber 590 may beconfigured to connect the sample entry port 595 to the conduit 630 inthe assembled cartridge 555.

In accordance with aspects of the present invention in which there aremultiple chips (e.g., a ground chip and a sensor chip), the cutaway 645may house one or more sensor chips 650 comprising at least one sensor655 (e.g., a PT, aPTT, or ACT micro-environment sensor), or a responsivesurface, together with an optional conductimetric sensor or sensors 660.The cutaway 665 may house a ground chip 670 comprising a groundelectrode 675 if needed as a return current path for an electrochemicalsensor, and may also house an optional conductimetric sensor. Inaccordance with aspects of the present invention in which there is onlya single chip, (e.g., a combined ground and sensor chip) the cutaway 665and the ground chip 670 may not be included with the cartridge 555.

In some embodiments, a metering means may be provided that comprises thesample holding chamber 590 bounded by the constriction or capillary stop640 and having along the sample holding chamber 590 length an air entrypoint 680 from the bladder comprising cavity 615. Air pressure exertedat the entry point 680 drives a metered volume of the sample past theconstriction or capillary stop 640. Therefore, a metered volume ofsample may be predetermined by a volume of the sample holding chamber590 between the air entry point 680 and the constriction or capillarystop 640. An amount of the sample corresponding to this volume may bedisplaced into the conduit 630 when the deformable member 605 isdisplaced. This arrangement may therefore provide a metering means fordelivering a metered amount of an unmetered sample into the variousdownstream conduits of the cartridge 555. The metering may beadvantageous in some embodiments if quantization of an analyte isrequired. Thus, an operator may be relieved of accurately measuring thevolume of the sample prior to measurement saving time, effort, andincreasing the accuracy and reproducibility.

In preferred embodiments, the invention is a process for using acartridge to determine diagnostic clotting times in a whole bloodsample. The process may include introducing an unmetered fluid sampleinto the sample holding chamber 590 of the cartridge 555 through thesample entry port 595 (as shown in FIG. 24). Capillary stop 640 preventspassage of the fluid sample into conduit 630 at this stage, and thesample holding chamber 590 is filled with the sample. Lid 585 is closedto prevent leakage of the fluid sample from the cartridge 555. Thecartridge 555 may then be inserted into the reading device or apparatus510, as shown in FIG. 23 and further disclosed in U.S. Pat. No.5,821,399, which is incorporated herein by reference in its entirety. Insome embodiments, insertion of the cartridge into the reading apparatus510 activates a mechanism, which punctures the fluid-containing packagelocated in the cavity 620 when the package is pressed against a spike(not shown). Fluid may thereby be expelled into one or more conduits(e.g., conduit 630) arriving in sequence at the sensor region.Thereafter, operation of a pump means (e.g., a pneumatic pump) appliespressure to the air-bladder comprised of cavity 615, forcing air througha conduit into the sample holding chamber 590 at the air entry point680. Capillary stop 640 delimits a metered portion of the original fluidsample. The metered portion of the sample is then expelled through thecapillary stop 640 by air pressure produced within air bladder comprisedof cavity 615. The sample passes into the conduit 630 and into contactwith the one or more reagents, the one or more substrates (e.g., animmobilized reagent-substrate-polymer layer), and/or the one or moresensors comprising one or more transducers and optionally the referenceelectrode located within the cutaway 665.

As also shown in FIG. 24, to promote (i) diffusion of the one or morereagents into the fluid sample, or diffusion of sample into the polymerlayer containing the reagents (depending on the specific embodiment),(ii) activation of the coagulation cascade by one of two pathways togenerate thrombin, (iii) diffusion of the active thrombin through theimmobilized substrate and/or reagent polymer layer, (iv) cleavage of thethrombin-cleavable peptide, (v) activation of the detectable moiety,and/or (vi) detection of the detectable moiety by the at least onetransducer, the fluid sample may be positioned within the conduit 630 tocontact the one or more reagents and/or substrates, the one or moreimmobilized polymer layers, and/or the one or more sensors, e.g.,micro-environment sensors, for a predetermined period of time.

Use of a cartridge is herein illustrated by a specific embodiment inwhich diagnostic clotting time is determined for a fluid sample, whichis introduced into the sample holding chamber of the cartridge followedby cartridge insertion into the cartridge reading device. The cartridgereading device makes electrical contact with the electrodes/sensorsthrough pads, and performs certain diagnostic tests. The diagnostictests determine whether fluid or sample is present in the conduits usingthe conductivity electrodes; determine whether electrical short circuitsare present in the electrodes; and ensure that the sensor and groundelectrodes are thermally equilibrated to, preferably, 37° C. prior tothe assay cycle.

In preferred embodiments, a metered portion of the fluid sample,preferably between 4 and 200 μL, more preferably between 4 and 20 μL,and most preferably 7 μL, may be used to carry out the assay, while asub-volume (between 0.1 and 3.5 uL) thereof may be used to contact theelectrodes/sensors. The fluid sample is positioned with respect to thesensor region such that a portion of the fluid sample is positioned overthe one or more reagents, the one or more substrates (e.g., immobilizedpolymer layers), and the one or more sensors comprising one or moretransducers and the ground electrode. After the predetermined period oftime, e.g., 0-10 seconds of oscillation in the upper or lower section of630 (or in any of the assay conduits in FIGS. 24, 25, 26, 31, 34, 36 forexample) the sample may move to a second conduit or area for subsequentmixing or interactions, or become static or become locked within theconduit(s) or cartridge prior to signal generation. One or moreconductivity sensors on the sensor chip may be used to control theseprocesses as discussed with respect to FIGS. 20, 21, 22A, and 22B.During subsequent splitting or diversion of the fluids, there may bepassage through pressure- or size-controlled elements. These aspects aredescribed in more detail later on.

During the time of contact between the sample and the sensors, (i) theamending reagents have time to diffuse into the fluid sample or thefluid sample has time to diffuse into the amending reagents (which couldbe immobilized in some embodiments) in order to promote activation ofthe coagulation cascade by one of two pathways to generate thrombin,(ii) the active thrombin has time to diffuse through a substrate layer,e.g., an immobilized substrate and/or reagent polymer layer, and cleavethe thrombin-cleavable peptide, and (iii) the activated detectablemoiety has time to be detected by the at least one transducer.

Fluidic Function and Configurations of Cartridges

In preferred embodiments, a disposable cartridge configuration isprovided for that enables two physically separated tests to be conductedsimultaneously or subsequently on a single whole blood sample within thesame disposable cartridge. The elements of the disposable cartridgeconfiguration include the use of passive fluidic features (e.g., valves,resistances, and fluidic locking elements) in addition to activemechanisms from the analyzer (e.g., a pump) to split the sample intoseparate conduits/regions such that each sample segment can subsequentlybe moved to a specific sensor. A number of separate configurations arediscussed herein which allow for maintaining the sample in a singlechannel, splitting the sample into separate fluidic conduits,controlling the fluid movement in each conduit to, for example, mixdried reagent and/or substrate into the sample segment, and/orsubsequently park (and lock) the sample over the sensors for analysis.However, it should be understood that various modifications,substitutions, omissions and changes of the configurations can be madewithout departing from the spirit and scope of the present invention.

In each embodiment below, the sample may be inserted into the inletchamber of the cartridge by a user. The cartridge is then closed andinserted into the analyzer. The diaphragm pump formed as an air bladderin the cartridge and a mechanical plunger in the analyzer (as discussedwith respect to FIG. 24) are used to move the sample throughout thecartridge.

The embodiments of the present invention discussed with respect to FIGS.25-29 are configured to split a single biological (e.g., whole blood)sample and allow independent mix control of at least two segments of thesample in two conduits where dried reagents and/or substrates specificto each test are located. In accordance with aspects of the presentinvention, the substrates may or may not be localized (e.g., inaccordance with some embodiments, may or may not be immobilized) overthe sensor. The reagents and/or substrates dissolved into the sampleremain within the conduit where they were formed, therefore eliminatingany potential cross-interference between the tests. This is an importantelement of multiplexing for any two tests where chemical or physicalinterference could be present (e.g., coagulation tests).

As shown in FIG. 25, some embodiments of the present invention pertainto a ground sensor first configuration for a cartridge 700 in which aconduit 705 splits at a junction 707 into a first conduit 710 and asecond conduit 715 prior to or upstream of a ground chip 720. The secondconduit 715 may comprise a constriction or capillary stop 725 and isconfigured to pass over a lower region of the sensor chip 730 comprisingat least one analyte detection electrode (as described with respect toFIGS. 15 and 16). The first conduit 710 is configured to pass over theground chip 720 (e.g., ground chip with reference sensor as describedwith respect to FIG. 17) and the upper region of the sensor chip 730comprising at least one analyte detection electrode (as described withrespect to FIGS. 15 and 16). The cartridge 700 may further comprise atleast one fluidic lock mechanism 735 (e.g., a membrane sponge valve, amicrochannel capillary, or a micro-array valve) positioned within thefirst conduit 710, and one or more conduits 740 (e.g., vents), whichlead from the first conduit 710 and the second conduit 715 to the cavity747. In this embodiment, the cavity 747 is configured as a waste chamber(as discussed with respect to FIG. 24). However, it should be understoodby those of skill in the art that the one or more conduits 740 may beconfigured to lead to a waste conduit (as discussed with respect to FIG.24).

FIG. 26 shows an alternative ground sensor first configuration for acartridge 750 in which the conduit 755 splits at a junction 760 into afirst conduit 765 and a second conduit 770 prior to or upstream of aground chip 775. The second conduit 770 may comprise a constriction orcapillary stop 780 and is configured to pass over a lower region of thesensor chip 785 comprising at least one analyte detection electrode (asdescribed with respect to FIGS. 15 and 16). The first conduit 765 isconfigured to pass over the ground chip 775 (e.g., ground chip withreference sensor as described with respect to FIG. 17) and the upperregion of the sensor chip 785 comprising at least one analyte detectionelectrode (as described with respect to FIGS. 15 and 16). The cartridge750 may further comprise at least one fluidic lock mechanism 790positioned within the first conduit 765. And one or more conduits 793and 795 (e.g., vents), which lead from the first conduit 765 and thesecond conduit 770 respectively to a waste conduit 797 (as discussedwith respect to FIG. 24).

As shown in FIG. 27, during operation of the cartridges 700 and 750, thefluid or biological sample is moved using the bi-directional diaphragmpump (as described with respect to FIG. 24) from the inlet or sampleholding chamber to the conduits 705/755. The biological sample is splitinto a first portion and a second portion at the junction 707/760 (e.g.,a T-junction). In preferred embodiments, the constriction or capillarystop 725/780 (e.g., a capillary burst valve or fluidicresistance/constriction) positioned within the second conduit 715/770causes the sample to preferentially fill the first conduit 710/765 andmove over the ground chip 720/775 and the at least one electrode (e.g.,an aPTT electrode) within the upper region of the sensor chip 730/785.The diaphragm pump can therefore move a first portion of the sample backand forth in the first conduit 710/765 to dissolve and mix the reagentand/or substrate into sample, while a second portion of the sample inthe second conduit 715/770 neither vacates the second conduit 715/770nor moves to the sensor chip 730/785. Once adequate mixing has beenachieved in the first conduit 710/765, the first portion of the sampleis pushed over the sensor chip 730/785 to the fluidic lock mechanism735/790 (e.g., a membrane “sponge valve” or a microchannel formed ineither the double-sided adhesive or one of the molded plasticcomponents), which provides pressure resistance and effectively locksthe first portion of the sample into the first conduit 710/765. Analysisin the first conduit 710/765 can thereafter begin. As the pressureresistance in the first conduit 710/765 increases significantly, theremaining second portion of the sample is forced through theconstriction or capillary stop 725/780 in the second conduit 715/770. Asimilar back and forth mixing process can then be applied to the secondconduit 715/770. Once the reagent and/or substrate is mixed into thesecond portion of the sample, the second portion of the sample can bepositioned over the sensor chip 730/785 and analysis in the secondconduit 715/770 can begin. For example, the second portion can be movedthrough the second conduit 715/770 over at least one electrode (e.g., aPT electrode) within the lower region of the sensor chip 730/785. Inaccordance with aspects of the present invention, the electrodes withinthe upper and lower region of the sensor chip 730/785 may be formed withor without immobilization of the reagent/substrate using one or more ofthe arrangements as discussed with respect to FIGS. 2, 3, 4, 7A, 7B, 7Cand 9.

As shown in FIG. 28, during additional operation of the cartridges 700and 750 the fluid or biological sample may be moved using thebi-directional diaphragm from the inlet or sample holding chamber to thesensor chip 730/785 (as described with respect to FIG. 24).Additionally, once the reagent and/or substrate is mixed into the secondportion of the sample, the second portion of the sample is pushed overthe sensor chip 730/785 to an additional fluidic lock mechanism 745/798(shown in FIGS. 25 and 26) (e.g., a membrane “sponge valve” or amicrochannel formed in either the double-sided adhesive or one of themolded plastic components), which provides pressure resistance andeffectively locks the second portion of the sample into the secondconduit 715/770. Once the second portion of the sample is locked inposition over the sensor chip 730/785, analysis in the second conduit715/770 can begin.

As shown in FIG. 29, during additional operation of the cartridges 700and 750 the fluid or biological sample may be moved using thebi-directional diaphragm from the inlet or sample holding chamber to theconduits 705/755. However, prior to the sample being moved to thejunction 707/760, the sample may be moved through an additional junction(e.g., T-junction) (not shown in FIGS. 25 and 26). The additionaljunction is configured to separate the first conduit 710/765 and thesecond conduit 715/770 from a relief conduit (e.g., a vent), which has aconstriction or capillary stop (e.g., a capillary burst valve or fluidicresistance/constriction) to divert flow of the sample through thejunction 707/760. Any residual pressure or movement of the sample willthen proceed into the relief conduit. In preferred embodiments, theadditional constriction or capillary stop in the relief conduit isdesigned such that it has a lower pressure resistance than the fluidiclocking features in the first conduit 710/765 and/or the second conduit715/770.

As shown in FIG. 30, the mixing within the conduits may be performed byoscillating the first and second portions of the sample over theelectrodes using the multi-conduit conductometric sensors to determineand maintain positioning of the sample as discussed herein. In someembodiments, the first portion of the biological sample is mixed firstto initiate a reaction between the first portion of the biologicalsample and the reagent and/or substrate within the first conduit priorto initiation of a reaction between the second portion of the biologicalsample and the reagent and/or substrate within the second conduit. Forexample, an aPTT test conventionally requires a longer test time thanthat of the PT test, and thus an aPTT test performed within the firstconduit could be started earlier than that of the PT test performedwithin the second conduit such that the tests are completed atapproximately the same time.

As should be understood, the ground sensor first design for thecartridges 700 and 750 advantageously provides for a single cartridgecapable of performing simultaneously or subsequently two independentassays (e.g., PT and aPTT) within two separate conduits. In embodimentsin which mixing is required or advantageous, the features of thecartridges 700 and 750 allow for independent mixing control within thefirst and second conduits without concern for cross-activation of thecascade pathways or other cross-electrode interference once the one ormore reagents have become exposed to the biological sample because thesensors are physically separate from one another via use of at least thefirst and second conduits.

The embodiments of the present invention discussed with respect to FIGS.31-35 are configured to split a single biological (e.g., whole blood)sample into two segments in two conduits where dried reagents and/orsubstrates specific to each test are located. The primary differencebetween the configurations discussed with respect to FIGS. 31-35 to thatof FIGS. 25-29 is that in the configurations of FIGS. 31-35 the sampleis pushed over the sensors and locked in place and the reagents and/orsubstrates located on the sensors are specifically designed to dissolveinto the sample by passive diffusion or remain within an immobilizationlayer. In the example of performing a coagulation analysis, theactivation of the cascade pathways and detection thereof occurs at ahigh concentration region in close proximity to the sensors. Thereagents and/or substrates dissolved into the sample or contained in theimmobilized layer remain within the conduit where they were printed,therefore eliminating any potential cross-interference between the tests(data not shown). This is an important element of multiplexing for anytwo tests where chemical or physical interference could be present(e.g., coagulation tests).

As shown in FIG. 31, some embodiments of the present invention pertainto an integrated ground sensor and split sensor conduit design for acartridge 800 in which a conduit 805 splits at a junction 810 into afirst conduit 815 and a second conduit 820 prior to or upstream of aground/sensor chip 825. The first conduit 815 is configured to pass overa first region of the ground/sensor chip 825 (as discussed with respectto FIG. 18) comprising a portion of the reference electrode and at leastone analyte detection electrode (e.g., an aPTT electrode). The secondconduit 820 comprises a constriction or capillary stop 822 and isconfigured to pass over a second region of the sensor chip 825 (asdiscussed with respect to FIG. 18) comprising another portion of thereference electrode and at least one other analyte detection electrode(e.g., a different analyte detection electrode, such as a PT electrode).In accordance with aspects of the present invention, the analytedetection electrodes within the first conduit 815 and the second conduit820 may be formed with or without immobilization of thereagent/substrate using one or more of the arrangements as discussedwith respect to FIGS. 2, 3, 4, 7A, 7B, 7C and 9.

The cartridge 800 may further comprise at least two fluidic barriermechanisms 830 and 835 (e.g., a fluidic lock mechanism, a capillarystop, or a fluidic constriction) positioned within the first conduit 815and the second conduit 820 respectively, and one or more conduits 840and 845 (e.g., vents), which lead from the first conduit 815 and thesecond conduit 820 respectively to a cavity 850. In this embodiment, thecavity 850 is configured as a waste chamber (as discussed with respectto FIG. 24). However, in alternative embodiments the one or moreconduits 840 and 845 may be configured to lead to a waste conduit (asdiscussed with respect to FIG. 24).

As shown in FIG. 32, during operation of the cartridge 800, the fluid orbiological sample is moved using the bi-directional diaphragm pump (asdescribed with respect to FIG. 24) from the inlet or sample holdingchamber to the conduit 805. The biological sample is split into a firstportion and a second portion at the junction 810 (e.g., a T-junction).In preferred embodiments, the constriction or capillary stop 822 (e.g.,a capillary burst valve or fluidic resistance/constriction) positionedwithin the second conduit 820 causes the sample to preferentially fillthe first conduit 815 and move over the first region of theground/sensor chip 825. The first portion of the sample is pushed overthe ground/sensor chip 825 to the fluidic barrier mechanism 830 (e.g., amembrane “sponge valve” or a microchannel formed in either thedouble-sided adhesive or one of the molded plastic components), whichprovides pressure resistance greater than that of the second conduit820, and effectively locks the first portion of the sample into thefirst conduit 815. Analysis in the first conduit 815 can thereafterbegin. As the pressure resistance in the first conduit 815 increasessignificantly, the remaining second portion of the sample is forcedthrough the constriction or capillary stop 822 in the second conduit820. Similarly, the second portion of the sample is pushed over theground/sensor chip 825 to the fluidic barrier mechanism 835 (e.g., amembrane “sponge valve” or a microchannel formed in either thedouble-sided adhesive or one of the molded plastic components), andeffectively locks the second portion of the sample into the secondconduit 820. Analysis in the second conduit 820 can thereafter begin.

As shown in FIG. 33, during additional operation of the cartridge 800the fluid or biological sample may be moved using the bi-directionaldiaphragm from the inlet or sample holding chamber to the conduit 805.However, prior to the sample being moved to the junction 810, the samplemay be moved through an additional junction 855 (e.g., T-junction)(shown in FIG. 31). The additional junction 855 (e.g., T-junction) isconfigured to separate the first conduit 810 and the second conduit 820from a relief conduit 860 (e.g., a vent), which has a constriction orcapillary stop (e.g., a capillary burst valve or fluidicresistance/constriction) to divert flow of the sample through thejunction 810. Any residual pressure or movement of the sample will thenproceed into the relief conduit 860. In preferred embodiments, theadditional constriction or capillary stop in the relief conduit isdesigned such that it has a lower pressure resistance than the fluidicbarrier mechanisms in the first and second conduits 815 and 820.

As should be understood by those of ordinary skill in the art, theintegrated ground sensor and split sensor conduit design for thecartridge 800 advantageously provides for a single cartridge capable ofperforming simultaneously or subsequently two independent assays (asshould be understood the assays may be different or the same, e.g., PTand aPTT, PT and PT, aPTT and aPTT, etc) within two separate conduitswithout the requirement of mixing the biological sample with the reagentand/or substrate. In accordance with aspects of this embodiment, thefeatures of the cartridge 800 allow for performing two separateanalytical tests within the first and second conduits 815 and 820without concern for cross-activation of the cascade pathways or othercross-electrode interference once the one or more reagents have becomeexposed to the biological sample because the electrodes are physicallyseparate from one another via use of at least the first and secondconduits 815 and 820. Moreover the integrated ground sensor design forthe cartridge 800 provides for a simpler more compact cartridge designthan that of the ground sensor first designs described above because thedesign eliminates the space requirement for a completely separate groundsensor and additional length of conduit necessary to move the biologicalsample to the separate ground sensor.

As shown in FIG. 34, alternative embodiments of the present inventionpertain to an integrated ground sensor and split conduit design for acartridge 870 comprising a conduit 875 with a junction 880 (e.g.,T-junction). The junction 880 is configured to separate a first sensorconduit 885 and a second conduit 890 from a relief conduit 895 (e.g., avent). The first sensor conduit 885 is configured to pass over a firstregion of a ground/sensor chip 900 (as discussed with respect to FIG.18) comprising a portion of a reference electrode and at least oneanalyte detection electrode (e.g., an aPTT electrode). The second sensorconduit 890 is configured to pass over a second region of the sensorchip 900 (as discussed with respect to FIG. 18) comprising anotherportion of the reference electrode and at least one different analytedetection electrode (e.g., a PT electrode). In accordance with aspectsof the present invention, the analyte detection electrodes within thefirst sensor conduit 885 and the second sensor conduit 890 may be formedwith or without immobilization of the substrate using one or more of thearrangements as discussed with respect to FIGS. 2, 3, 4, 7A, 7B, 7C and9.

The cartridge 870 may further comprise at least two fluidic barriermechanisms 905 and 910 positioned within the first sensor conduit 885and the second sensor conduit 890 respectively, and one or more conduits915 and 920 (e.g., vents), which lead from the first sensor conduit 885and the second conduit 890 respectively to a cavity 925. In thisembodiment, the cavity 925 is configured as a waste chamber (asdiscussed with respect to FIG. 24). However, in alternative embodimentsthe one or more conduits 915 and 920 may be configured to lead to awaste conduit (as discussed with respect to FIG. 24).

As shown in FIG. 35, during operation of the cartridges 870, the fluidor biological sample is moved using the bi-directional diaphragm pump(as described with respect to FIG. 24) from the inlet or sample holdingchamber to the conduit 875 and through the junction 880. The reliefconduit 895 (e.g., a vent) has a constriction or capillary stop (e.g., acapillary burst valve or fluidic resistance/constriction) to divert flowof the sample to the first sensor conduit 885 and the second sensorconduit 890. A first portion of the sample is pushed over theground/sensor chip 900 to the fluidic barrier mechanism 905 (e.g., amembrane “sponge valve” or a microchannel formed in either thedouble-sided adhesive or one of the molded plastic components), whicheffectively locks the first portion of the sample into the first sensorconduit 885. Analysis in the first conduit 885 can thereafter begin. Asecond portion of the sample is pushed over the ground/sensor chip 900to the fluidic barrier mechanism 910 (e.g., a membrane “sponge valve” ora microchannel formed in either the double-sided adhesive or one of themolded plastic components), which effectively locks the second portionof the sample into the second sensor conduit 890. Analysis in the secondconduit 890 can thereafter begin. Any residual pressure or movement ofthe sample will then proceed into the relief conduit 895. In preferredembodiments, the constriction or capillary stop in the relief conduit isdesigned such that it has a lower pressure resistance than the fluidicbarrier mechanisms in the first and second sensor conduits 885 and 890.

As should be understood, the alternative integrated ground sensor andsplit conduit design for the cartridge 870 advantageously provides for asingle cartridge capable of performing simultaneously or subsequentlytwo independent assays (as should be understood the assays may bedifferent or the same, e.g., PT and aPTT, PT and PT, aPTT and aPTT,etc.) within two separate sensor conduits without the requirement ofmixing the biological sample with the reagent and/or substrate. Inaccordance with aspects of this embodiment, the features of thecartridge 870 allow for performing two separate analytical tests withinthe first and second sensor conduits 885 and 890 without concern forcross-activation of the cascade pathways or other cross-electrodeinterference once the one or more reagents have become exposed to thebiological sample because the electrodes are physically separate fromone another via use of at least the first and second sensor conduits 885and 890. Moreover. the integrated ground sensor design for the cartridge870 provides for a simpler more compact cartridge design than that ofthe ground sensor first designs described above because the designeliminates the space requirement for a completely separate ground sensorand additional length of conduit necessary to move the biological sampleto the separate ground sensor. In addition, the sample volume requiredto cover the complete sensor circuit is significantly reduced.

The embodiments of the present invention discussed with respect to FIGS.36-38 are configured to maintain a single biological (e.g., whole blood)sample in a single conduit where dried reagents and/or substratesspecific to each test are located. The primary difference between theconfigurations discussed with respect to FIGS. 36-38 to that of FIGS.27-29 is that in the configurations of FIGS. 36-38 the sample ismaintained in a single conduit and pushed over the sensors in series andlocked or held in place and the reagents and/or substrates located onthe sensors are specifically designed to dissolve into the sample bypassive diffusion or remain within an immobilization layer. In theexample of performing a coagulation analysis, the activation of thecascade pathways and detection thereof occurs at a high concentrationregion in close proximity to the sensors. The reagents and/or substratesdissolved into the sample remain near the sensors where they wereprinted, therefore eliminating any potential cross-interference betweenthe tests. This is an important element of multiplexing for any twotests where chemical or physical interference could be present (e.g.,coagulation tests).

As shown in FIG. 36, some embodiments of the present invention pertainto an integrated ground sensor and single conduit design for a cartridge930 comprising a conduit 935 with a junction 940 (e.g., T-junction). Thejunction 940 is configured to separate a single conduit sensor conduit945 from a relief conduit 950 (e.g., a vent). The single sensor conduit945 is configured to pass over a first region of a ground/sensor chip955 (as discussed with respect to FIG. 18) comprising at least oneanalyte detection electrode (e.g., an aPTT electrode), a second regioncomprising at least a portion of the reference electrode, and a thirdregion comprising at least one same or different analyte detectionelectrode (e.g., a PT electrode). In accordance with aspects of thepresent invention, the analyte detection electrodes within the singleconduit sensor conduit 945 are formed with immobilization of thereagent/substrate using one or more of the arrangements as discussedwith respect to FIGS. 4, 7A, 7B, 7C and 9.

The cartridge 930 may further comprise a fluidic barrier mechanism 960positioned within the sensor conduit 945, and a conduit 965 (e.g.,vents) that leads from the sensor conduit 945 to a cavity 970. In thisembodiment, the cavity 970 is configured as a waste chamber (asdiscussed with respect to FIG. 24). However, in alternative embodimentsthe conduit 965 may be configured to lead to a waste conduit (asdiscussed with respect to FIG. 24).

As shown in FIGS. 36 and 37, during operation of the cartridge 930, thefluid or biological sample is moved using the bi-directional diaphragmpump (as described with respect to FIG. 24) from the inlet or sampleholding chamber to the conduit 935 and through the junction 940. Therelief conduit 950 (e.g., a vent) has a constriction or capillary stop(e.g., a capillary burst valve or fluidic resistance/constriction) todivert flow of the sample to the sensor conduit 945. The sample ispushed over the ground/sensor chip 955 to the fluidic barrier mechanism960 (e.g., a membrane “sponge valve” or a microchannel formed in eitherthe double-sided adhesive or one of the molded plastic components),which effectively locks the sample into the sensor conduit 945. Analysisin the sensor conduit 945 can thereafter begin. Any residual pressure ormovement of the sample will then proceed into the relief conduit 950. Inpreferred embodiments, the constriction or capillary stop in the reliefconduit 950 is designed such that it has a lower pressure resistancethan the fluidic barrier mechanisms in the sensor conduit 945.

As shown in FIGS. 36 and 38, during alternative operation of thecartridge 930, the fluid or biological sample may be moved using thebi-directional diaphragm from the inlet or sample holding chamber to theconduit 935. The sample is thereafter pushed over the ground/sensor chip955 to the fluidic barrier mechanism 960 (e.g., a membrane “spongevalve” or a microchannel formed in either the double-sided adhesive orone of the molded plastic components), which effectively locks thesample into the sensor conduit 945. Analysis in the sensor conduit 945can thereafter begin.

As should be understood, the integrated ground sensor and single conduitdesign for the cartridge 930 advantageously provides for a singlecartridge capable of performing simultaneously or subsequently twoindependent assays (e.g., as should be understood the assays may bedifferent or the same, e.g., PT and aPTT, PT and PT, aPTT and aPTT, etc)within a single conduit without the requirement of mixing the biologicalsample with the reagent and/or substrate. In accordance with aspects ofthis embodiment, the features of the cartridge 930 allow for performingtwo separate analytical tests within the sensor conduit 945 withoutconcern for cross-activation of the cascade pathways or othercross-electrode interference once the one or more reagents have becomeexposed to the biological sample because the analyte detectionelectrodes are micro-environment sensors with a localized (e.g.,immobilized) reagent/substrate formed using one or more of thearrangements as discussed herein in detail. Moreover the integratedground sensor design for the cartridge 930 provides for a simpler morecompact cartridge design than that of the ground sensor first designsdescribed above because the design eliminates the space requirement fora completely separate ground sensor and additional length of conduitnecessary to move the biological sample to the separate ground sensor.In addition, the sample volume required to cover the complete sensorcircuit is significantly reduced.

The embodiments of the present invention discussed with respect to FIG.39 are configured to split a single biological (e.g., whole blood)sample and allow for multiple physically separated tests (e.g.,coagulation tests and analytical chemistry tests) to be conductedsimultaneously or subsequently on a single whole blood sample within thesame disposable cartridge. In some embodiments, independent mix controlis provided for at least two segments of the sample in two conduitswhere dried reagents and/or substrates specific to each test may belocated. In accordance with aspects of the present invention, thesubstrates may or may not be localized (e.g., immobilized) over thesensors. The reagents and/or substrates dissolved into the sample remainwithin the conduit and close to the sensor where they were printed ordeposited, therefore eliminating any potential cross-interferencebetween the tests. This is an important element of multiplexing for anytwo tests where chemical or physical interference could be present(e.g., coagulation tests and analytical chemistry tests).

As shown in FIG. 39, some embodiments of the present invention pertainto a multiple sensor configuration for a cartridge 1000 comprising aconduit 1005 with a junction 1010 (e.g., T-junction). The junction 1010may comprise a first constriction or capillary stop 1015 and isconfigured to separate a first sensor conduit 1020 from an ancillaryconduit 1025. The first sensor conduit 1020 is configured to pass overat least a portion of a sensor chip 1030 comprising at least one analytedetection electrode (e.g., a PT electrode). In accordance with aspectsof the present invention, the analyte detection electrode within thefirst sensor conduit 1020 may be formed with or without localization(e.g., immobilization) of a reagent/substrate using one or more of thearrangements as discussed with respect to FIGS. 2, 3, 4, 7A, 7B, 7C and9. A second constriction or capillary stop 1035 may be positioned withinthe first sensor conduit 1020, and a conduit 1040 (e.g., vents) may beconfigured to lead from the first sensor conduit 1020 to a waste conduit1045 (as discussed with respect to FIG. 24).

The cartridge 1000 may further comprise a second sensor conduit 1050that connects a cavity 1055 with the waste conduit 1045. The ancillaryconduit 1025 connects to the second sensor conduit 1050 at a junction1060. The junction 1060 may comprise a third constriction or capillarystop 1065. In some embodiments, the first constriction or capillary stop1015 and the third constriction or capillary stop 1065 are configuredlarger (e.g., larger in width) than the second constriction or capillarystop 1035 to allow for control of the sample as discussed hereafter indetail. The second sensor conduit 1050 is configured to pass over atleast a portion of each of one or more sensor chips 1070 comprising atleast one analyte detection electrode (e.g., a sodium or chlorideelectrode). The one or more sensor chips 1070 may be configured toperform any number of assays, including electrolytes, generalchemistries, blood gases and hematology (See, for example, U.S. Pat.Nos. 7,419,821, 6,379,883, 5,514,253, 5,200,051, and 5,096,669, whichare incorporated herein by reference in their entireties). For example,the one or more sensor chips 1070 may be configured to perform anynumber of assays capable of detecting one or more analytes selected fromthe group consisting of oxygen partial pressure, carbon dioxide partialpressure, total carbon dioxide, pH, potassium, sodium, chloride,glucose, BUN, creatinine, lactate, magnesium, hematocrit, ionizedcalcium, troponin I, troponin T, CKMB, procalcitonin, bHCG, HCG,NTproBNP, proBNP, BNP, myoglobin, parathyroid hormone, d-dimer, NGAL,galectin-3, and/or PSA, among other analytes. In embodiments in which asubstrate is utilized for performing the assay, the at least one analytedetection electrode within the second sensor conduit 1050 may be formedwith or without localization (e.g., immobilization) of areagent/substrate using one or more of the arrangements as discussedwith respect to FIGS. 2, 3, 4, 7A, 7B, 7C and 9. In other embodimentsthat do not utilize a substrate, the at least one analyte detectionelectrode within the second sensor conduit 1050 may be formed withoutany substrate.

As shown in FIG. 39, during operation of the cartridge 1000, deformationof a gasket by the analyzer may transmit pressure onto afluid-containing foil pack filled with a fluid, e.g., approximately 130μL of analysis/wash solution, control fluid, or calibrant fluid, locatedin the cavity 1055, rupturing the foil pack, and expelling fluid intothe second sensor conduit 1050 and past the third constriction orcapillary stop 1065 for subsequent use in sample analysis (as discussedwith respect to FIG. 24). In some embodiments, the foil pack is acalibrant package (CALPAK) that contains a calibrant solution. Thetypical sequence of events includes the CALPAK being ruptured and thenthe calibration solution passing over the one or more sensor chips 1070to wet up the one or more sensor chips 1070. Thereafter, the fluid orbiological sample is moved using the bi-directional diaphragm pump (asdescribed with respect to FIG. 24) from the inlet or sample holdingchamber to the conduit 1005. The biological sample is split into a firstportion and a second portion at the junction 1010 (e.g., a T-junction).In preferred embodiments, the first constriction or capillary stop 1015(e.g., a capillary burst valve or fluidic resistance/constriction)positioned within the ancillary conduit 1025 causes the first portion ofthe sample to preferentially fill the first sensor conduit 1020 and moveover the sensor chip 1030 and the at least one electrode (e.g., a PTelectrode). The first portion of the sample within the first sensorconduit 1020 stops at the second constriction or capillary stop 1035,which causes the second portion of the sample to push through the firstconstriction or capillary stop 1015 and the third constriction orcapillary stop 1065. The second portion of the sample fills the secondsensor conduit 1050 and moves over one or more sensor chips 1070 and theat least one analyte detection electrode (e.g., a sodium or chlorideelectrode). In some embodiments, the second portion of the sample may beconfigured to mix with the analysis/wash solution, control fluid, orcalibrant fluid present within the second sensor conduit 1050. To enablemixing of the fluid segments, features, including retention structuressuch as post arrays, conduit cut-outs, grooves, or dimples, can bedesigned into second sensor conduit 1050 to retain the analysis/washsolution, control fluid, or calibrant fluid. Alternatively, mixing ofthe two fluid segments can be accomplished by merging the two streams offluid from conduit 1025 and the foil-pack 1055 at junction 1060. Inother embodiments, the analysis/wash solution, control fluid, orcalibrant may have been pumped through the second sensor conduit 1050 tothe waste conduit 1045 such that the second portion of the sample doesnot mix with the analysis/wash solution, control fluid, or calibrant.Furthermore, to minimize carry-over of the analysis/wash solution,control fluid, or calibrant, the cartridge may be designed in such a wayas to introduce an air segment between the first fluid and the secondportion of the sample. The volume of ancillary conduit 1025 shoulddetermine the size of the air gap between the fluid segments.

As should be understood, the multiple sensor configurations for thecartridge 1000 advantageously provide for a single cartridge capable ofperforming simultaneously or subsequently two independent assays (e.g.,PT and an analytical chemistry assay, PT and PT, aPTT and PT, aPTT andaPTT, etc.) within two separate conduits. In embodiments in which mixingis required or advantageous, the features of the cartridge 1000 allowfor independent mixing control within the first and second conduitswithout concern for cross-activation or other cross-electrodeinterference once the one or more reagents have become exposed to thebiological sample because the sensors are physically separate from oneanother via use of at least the first and second sensor conduits.

As shown in FIG. 40 where the x-axis is time/seconds and the y-axis iscurrent/pA, embodiments pertaining to independent mix control of asample in two separate conduits of a cartridge (as shown in FIGS. 25-29and 39) and fully mixing the reagent/substrate (i.e., no localizedmicro-environment) are capable of achieving clotting results (for e.g.,PT times indicated by the vertical line marked on each of the clotcurves) using multiple sample types (whole blood represented by responsecurves 1071, two levels of factor depleted blood represented by responsecurves 1072 and 1073, one level of control plasma represented byresponse curves 1075). As shown in FIG. 41A, embodiments of the presentinvention pertaining to no mixing of a sample over the immobilized orlocalized reagents/substrates in one or more conduits of a cartridge (asshown in FIGS. 30-38) are also capable of achieving clotting resultsusing samples (whole blood represented by response curves 1076, factordepleted blood represented by response curves 1077, and two levels ofcontrol plasma represented by response curves 1078 and 1079) similar tothose used in FIG. 40. As shown in FIG. 41B (whole blood represented byresponse curves 1076, factor depleted blood represented by responsecurves 1077, and two levels of control plasma represented by responsecurves 1078 and 1079), the sensor responses for the embodimentspertaining to mixing of the sample over the immobilized or localizedreagents/substrates are also achievable, however resolution of thefactor depleted and extended plasma control samples is not as clear asin the immobilized no mixing embodiment (as shown in FIG. 41A). Further,the preference is that the clotting times of whole blood (curves 1071and 1076) be very close to the normal plasma control clotting times(curves 1072 and 1078); the no mixed, immobilized/localized embodimentsreflect this the best, further giving credence to their improvement overthe system of FIG. 40. Although no-mixed or mixed versions of theembodiments are possible, the current (pA) generation is higher andvariability (pA and clotting time) is lower in the no-mixed version (asshown in FIG. 41A) of the embodiments. These unexpected and moreconsistent results are directly attributable to the use of themicro-environment sensors as described herein with respect to the nomixing embodiments. The no mixing embodiment in combination with thelocalized (e.g., immobilized) reagent/substrate print represents asystem improvement in the activation/propagation of the coagulationsignal in a higher reagent/substrate to lower sample volume ratio, thusproducing a faster assay/sensor response time. In addition, localization(e.g., immobilization) of the reagent/substrate print directly over thesensor results in the immediate turnover (oxidation) of the diffusedsubstrate leaving group once generated by active thrombin. Thisultimately leads to greater (and more reproducible) current signal beinggenerated directly at the amperometric assay sensor. Finally, thecombination of quick thrombin response (evident from the quick rise incurves in FIG. 41A as compared to the slower rise in FIG. 41B) withhigher currents and higher reproducibility produces response curveswhich are more easily and reproducibly analyzed, thus yielding animproved assay (the current embodiment, represented in FIG. 41A) overeither of those in FIG. 40 or 41B.

Ground Chip Elimination and Cartridge Identification

In preferred embodiments, the ground chip may be incorporated orintegrated into the sensor chip as described in detail herein. A typicalground chip (as described with respect to FIG. 17) may include a groundelectrode serving as a return of the sensor chip and four contact pinsfor cartridge identification (see, e.g., U.S. Pat. No. 7,419,821, whichis incorporated herein in its entirety). Accordingly, integration of theground chip into the sensor chip encompasses moving these two functionsinto the sensor chip. The advantages of integrating the ground chip withthe sensor chip include (i) a simplified manufacturing process sincethere is one component less to deal with during wafer fabrication,metallization, dicing, and cartridge assembly, (ii) reduced cost, and(iii) reduced sample volume since the sensor conduit can be shortened asshown in a comparison between at least FIGS. 25 and 31.

As shown in FIG. 42, the separate ground chip and sensor chiparrangement 1080 (e.g., the arrangement shown in FIG. 17) typicallyfunctions by using a detector 1082 connected with the ground pin 1085 onthe ground chip 1090 and an amperometric pin 1095 on the sensor chip1100 to detect a difference in current between the reference electrode1105 and the analyte detection electrode 1110 (e.g., an amperometricelectrode). In order to impart the reference electrode functionalityinto a single sensor chip arrangement 1115 (e.g., the arrangement shownin FIG. 18) the reference electrode 1105 may be integrated with thesensor chip 1100 by connecting the reference electrode 1105 with theconductometric low pin 1120. An electronic switch 1125 may beimplemented in the analyzer, which is configured to connect the groundpin 1085 and the conductometric low pin 1120. Accordingly, the referenceelectrode 1105 can be essentially moved from the ground chip 1090 to thesensor chip 1100.

In order to impart the cartridge identification functionality into asingle sensor chip arrangement (e.g., the arrangement shown in FIG. 18),an additional mechanism or means may be included in the arrangement forcartridge identification. As shown in FIG. 42 for a single electrodearrangement 1130, for example, an aPTT only analyte detection electrode1110, a resistor 1135 can be implemented between an unused amperometricpin 1140 and the conductometric low pin 1120. The analyte detectionelectrode 1110 may be connected to the amperometric pin 1095, theresistor 1135 may be connected to the unused amperometric pin 1140 andthe conductometric low pin 1120, and the reference electrode 1105 may beconnected to the conductometric low pin 1120. The resistance of theresistor 1135 may be measured by a detector 1145 (e.g., processor) byapplying a small voltage, e.g., 1 mV, between the unused amperometricpin 1140 and the conductometric low pin 1120, subsequent to (e.g.,immediately after) the cartridge being inserted into the analyzer. Thevalue of the measured resistance can then be used for cartridgeidentification. For example, each cartridge type (e.g., i-STAT®cartridges EC8+, CG8+, EG7+, CHEM8+, etc.) may be associated with acertain resistance or resistance range such that a measured resistanceof the cartridge may be used to identify the type of cartridge using alook-up table.

In some embodiments, the resistor 1135 may be comprised of a metal wire,preferably a gold wire manufactured at a same time as the contact padsand sensor electrode. The gold wire may be as small as 5 μm wide and 0.1μm thick, which forms an area of 0.5 μm². As the resistivity of gold is2.44 μΩ-cm, or 0.0244 Ω-μm, a 1000 μm long gold wire will have aresistance of 0.0244 Ω-μm*1000 μm/0.5 μm²=48.8Ω. After the cartridge isinserted into the analyzer, a small voltage, e.g., 0.5 mV may be appliedand a current of around 10 uA may be generated and detected by theanalyzer. To minimize the power consumption, optionally the gold wirecould be longer, the applied voltage could be lower, or the time for theapplication of the voltage could be shorter.

In an alternative embodiment, the single electrode arrangement 1130 mayinclude a PT only analyte detection electrode 1110 rather than an aPTTonly analyte detection electrode 1110. In accordance with this aspect ofthe present invention, the length of the gold wire may be increased toabout 10 cm, which increases the resistance of the gold wire to around5000Ω, in order to distinguish identification of the PT cartridge fromthat of the aPTT cartridge.

In other embodiments, the resistor may be implemented between theamperometric pin 1095 and the conductometric low pin 1120. As should beunderstood by those of ordinary skill in the art, the concept of using aresistor to identify the type of cartridge may be implemented in any ofthe sensor/cartridge arrangements described herein. Moreover differentvalues for the resistor can be obtained by varying the geometries of thewire or using varied materials for the wire (e.g., using TiW rather thangold), which can then be used for identifying different cartridgeswithout departing from the spirit and scope of the present invention.

In order to impart the cartridge identification functionality into amultiple sensor chip arrangement (e.g., the arrangement shown in FIG.18), an additional mechanism or means may be included in the arrangementfor cartridge identification. As shown in FIG. 43 for a multiple (e.g.,two) electrode arrangement 1150, for example, a PT analyte detectionelectrode 1155 and an aPTT analyte detection electrode 1160, a resistor1165 can be implemented between an unused amperometric pin 1170 and theconductometric low pin 1175. The PT analyte detection electrode 1155 maybe connected to the amperometric pin 1180, the aPTT analyte detectionelectrode 1160 may be connected to the amperometric pin 1185, theresistor 1165 may be connected to the unused amperometric pin 1170 andthe conductometric low pin 1175, and the reference electrode 1190 may beconnected to the conductometric low pin 1175. The resistance of theresistor 1165 may be measured by a detector 1195 by applying a smallvoltage, e.g., 1 mV, between the unused amperometric pin 1170 and theconductometric low pin 1175, subsequent to (e.g., immediately after) thecartridge being inserted into the analyzer. The value of the measuredresistance can then be used for cartridge identification. For example,each cartridge type (e.g., i-STAT® cartridges EC8+, CG8+, EG7+, CHEM8+,etc.) may be associated with a certain resistance or resistance rangesuch that a measured resistance of the cartridge may be used to identifythe type of cartridge using a look-up table.

As discussed above, the resistor 1165 may be comprised of a metal wire,preferably a gold wire manufactured at a same time as the contact padsand sensor electrode. The gold wire may be as small as 5 μm wide and 0.1μm thick, which forms an area of 0.5 μm². As the resistivity of gold is2.44 μΩ-cm, or 0.0244 Ω-μm, a 1000 μm long gold wire will have aresistance of 0.0244 Ω-μm*1000 μm/0.5 μm²=48.8Ω. After the cartridge isinserted into the analyzer, a small voltage, e.g., 0.5 mV may be appliedand a current of around 10 uA may be generated and detected by theanalyzer. To minimize the power consumption, optionally the gold wirecould be longer, the applied voltage could be lower, or the time for theapplication of the voltage could be shorter.

While the invention has been described in terms of various preferredembodiments, those skilled in the art will recognize that variousmodifications, substitutions, omissions and changes can be made withoutdeparting from the spirit of the present invention. It is intended thatthe scope of the present invention be limited solely by the scope of thefollowing claims. In addition, it should be appreciated by those skilledin the art that a plurality of the various embodiments of the invention,as described above, may be coupled with one another and incorporatedinto a single reader device.

We claim:
 1. An integrated circuit chip comprising: a micro-environmentprothrombin time (PT) sensor; a micro-environment activated partialthromboplastin time (aPTT) sensor or an micro-environment activatedclotting time (ACT) sensor, wherein the micro-environment PT sensor ison a same vertical plane as the micro-environment aPTT sensor or themicro-environment ACT sensor; a first temporary electrical connectorconnected via a wiring to the micro-environment PT sensor; and a secondtemporary electrical connector connected via another wiring to themicro-environment aPTT sensor or the micro-environment ACT sensor. 2.The integrated circuit chip of claim 1, wherein the micro-environment PTsensor comprises at least one reagent for inducing extrinsic coagulationpathway.
 3. The integrated circuit chip of claim 1, wherein themicro-environment aPTT sensor or the micro-environment ACT sensorcomprises at least one reagent for inducing intrinsic coagulationpathway.
 4. The integrated circuit chip of claim 1, wherein themicro-environment PT sensor is an amperometric PT sensor, wherein themicro-environment aPTT sensor is an amperometric aPTT sensor, whereinthe micro-environment ACT sensor is an amperometric ACT sensor, andwherein the first temporary electrical connector is an amperometricconnector, and the second temporary electrical connector is anamperometric connector.
 5. The integrated circuit chip of claim 1,further comprising a counter-reference electrode.
 6. The integratedcircuit chip of claim 5, wherein the counter-reference electrodecompletes a circuit on the integrated circuit chip, and the circuitincludes the micro-environment PT sensor and the micro-environment aPTTsensor or the micro-environment ACT sensor.
 7. The integrated circuitchip of claim 1, further comprising at least two conductometric sensors.8. The integrated circuit chip of claim 1, wherein the micro-environmentPT sensor comprises at least one transducer coated with a polymer layer,and the polymer layer comprises a thrombin-cleavable peptide with adetectable moiety.
 9. The integrated circuit chip of claim 1, whereinthe micro-environment aPTT sensor or the micro-environment ACT sensorcomprises at least one transducer coated with a polymer layer, and thepolymer layer comprises a thrombin-cleavable peptide with a detectablemoiety.
 10. An integrated circuit chip comprising: a test sensorcomprising at least one transducer coated with a polymer layer, whereinthe polymer layer is a porous support layer that comprises immobilizedtherein a thrombin-cleavable peptide; and a first temporary electricalconnector connected via a wiring to the test sensor.
 11. The integratedcircuit chip of claim 10, wherein the test sensor is a prothrombin time(PT) sensor.
 12. The integrated circuit chip of claim 11, wherein the PTsensor comprises at least one reagent for inducing extrinsic coagulationpathway.
 13. The integrated circuit chip of claim 10, wherein the testsensor is an activated partial thromboplastin time (aPTT) sensor. 14.The integrated circuit chip of claim 13, wherein the aPTT sensorcomprises at least one reagent for inducing intrinsic coagulationpathway.
 15. The integrated circuit chip of claim 10, wherein the testsensor is activated clotting time (ACT) sensor.
 16. The integratedcircuit chip of claim 15, wherein the ACT sensor comprises at least onereagent for inducing intrinsic coagulation pathway.
 17. The integratedcircuit chip of claim 10, wherein the thrombin-cleavable peptidecomprises a detectable moiety linked by a thrombin-cleavable amide bondto a polypeptide sequence that is non-reactive with blood proteasesother than thrombin.
 18. The integrated circuit chip of claim 10,wherein the test sensor further comprises a reagent for a bloodcoagulation assay, and wherein the reagent is formed: (i) as a layerover the polymer layer, (ii) within the polymer layer, or (iii) in alocation adjacent to the polymer layer.
 19. The integrated circuit chipof claim 10, wherein the test sensor is an electrochemical sensor or anoptical sensor.
 20. The integrated circuit chip of claim 19, wherein theelectrochemical sensor is an amperometric sensor or a potentiometricsensor.